Epidermal Photonic Systems and Methods

ABSTRACT

The invention provides systems and methods for tissue-mounted photonics. Devices of some embodiments implement photonic sensing and actuation in flexible and/stretchable device architectures compatible with achieving long term, mechanically robust conformal integration with a range of tissue classes, including in vivo biometric sensing for internal and external tissues. Tissue-mounted photonic systems of some embodiments include colorimetric, fluorometric and/or spectroscopic photonics sensors provided in pixelated array formats on soft, elastomeric substrates to achieve spatially and/or or temporally resolved sensing of tissue and/or environmental properties, while minimize adverse physical effects to the tissue. Tissue-mounted photonic systems of some embodiments enable flexible passive or active optical sensing modalities, including sensing compatible with optical readout using a mobile electronic devices such as a mobile phone or tablet computer.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. patent application Ser. No.15/501,364, filed Feb. 2, 2017, which is a U.S. National StageApplication under 35 U.S.C. § 371 of International Application No.PCT/US2015/044573, filed Aug. 11, 2015, which claims the benefit of andpriority to U.S. Provisional Patent Application No. 62/035,823, filedAug. 11, 2014, U.S. Provisional Patent Application No. 62/035,866, filedAug. 11, 2014, and U.S. Provisional Patent Application No. 62/142,877,filed Apr. 3, 2015, each of which is hereby incorporated by reference inits entirety to the extent not inconsistent herewith.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT

This invention was made with governmental support under GrantN00014-10-1-0989 awarded by the Office of Naval Research. The governmenthas certain rights in the invention.

BACKGROUND OF INVENTION

Wearable electronics and photonics are a class of systems with potentialto broadly impact a range of technologies, industries and consumerproducts. Advances in wearable systems are driven, in part, bydevelopment of new materials and device architectures providing for newfunctionalities implemented using device form factors compatible withthe body. Wearable consumer products are available, for example, thatexploit small and portable electronic and/or photonic systems providedin body mounted form factors, such as systems building off ofconventional body worn devices such as eye glasses, wrist bands, footware, etc. New device platforms are also under development to extend therange of wearable technology applications including smart textiles andstretchable/flexible electronic systems incorporating advancedelectronic and photonic functionality in spatially complaint formfactors compatible with low power operation, wireless communication andnovel integration schemes for interfacing with the body. [see, e.g., Kimet al., Annu. Rev. Biomed. Eng. 2012. 14; 113-128; Windmiller, et al.,Electroanalysis; 2013, 25, 1, 29-46; Zeng et al., Adv. Mater., 2014, 26,5310-5336; Ahn et al., J Phys. D: Appl. Phys., 2012, 45, 103001].

Tissue mounted systems represents one class of wearable systemssupporting diverse applications in healthcare, sensing, motionrecognition and communication. Recent advances in epidermal electronics,for example, provide a class of skin-mounted electronic systems providedin physical formats enabling mechanically robust and physically intimatecontact with the skin. Certain classes of epidermal electronic systemshave been developed, for example, combining high performance stretchableand/or ultrathin functional materials with soft elastic substratesimplemented in device geometries useful for establishing and maintainingconformal contact with the soft, curvilinear and time varying surface ofthe skin. [see, e.g., US Publication No. 2013/0041235]W.-H. Yeo, Y.-S.Kim, J. Lee, A. Ameen, L. Shi, M. Li, S. Wang, R. Ma, S. H. Jin, Z.Kang, Y. Huang and J. A. Rogers, “Multifunctional Epidermal ElectronicsPrinted Directly Onto the Skin,” Advanced Materials 25, 2773-2778(2013). Important to adoption of the emerging class of epidermalelectronic systems is the continued development devices supporting awide range of applications for this technology including for personalhealthcare assessment and clinical medicine.

It will be appreciated from the foregoing that tissue mounted systemsare needed to support the rapidly emerging applications in wearableelectronics. New epidermal systems are needed, for example, providingnew sensing, readout and analysis modalities to support diversetechnology applications in physiological and environmental sensing.

SUMMARY OF THE INVENTION

The invention provides systems and methods for tissue-mounted photonics.Devices of some embodiments implement photonic sensing and actuation inflexible and/stretchable device architectures compatible with achievinglong term, mechanically robust conformal integration with a range oftissue classes, including in vivo biometric sensing for internal andexternal tissues. Tissue-mounted photonic systems of some embodimentsinclude colorimetric, fluorometric and/or spectroscopic photonicsstructures provided in pixelated array formats on soft, elastomericsubstrates to achieve spatially and/or or temporally resolved sensing oftissue and/or environmental properties, while minimize adverse physicaleffects to the tissue. Tissue-mounted photonic systems of someembodiments enable robust and convenient optical sensing modalities,including sensing compatible with optical readout using a mobileelectronic devices such as using the camera and processor of a mobilephone or tablet computer. Tissue-mounted photonic systems of someembodiments have a low effective modulus and small thickness providingmechanical properties compatible with a range of deployment modes suchas direct adhesion on the surface of a tissue and deployment usingadhesives or intermediate bonding structures.

In one aspect, the invention provides a photonic device for interfacingwith a tissue, the device comprising: (i) a flexible or stretchablesubstrate; and (ii) one or more photonic structures supported by theflexible or stretchable substrate for generating a photonic responsecorresponding to one or more tissue parameters or environmentalparameters; wherein the flexible or stretchable substrate and the one ormore photonic structures provide a net bending stiffness (and/or Young'smodulus) such that the device is capable of establishing conformalcontact with a surface of the tissue. In an embodiment, the device isfor spatial and/or temporally characterizing tissue parameters orenvironmental parameters, for example, in connection withcharacterization of physiological, chemical and or environmentproperties of the tissue at, or below, the surface of the tissue and/orcorresponding to materials derived from the tissue, e.g., biofluids. Inan embodiment, for example, the device is for sensing or actuating thetissue. In an embodiment, for example, the device is for the device isfor sensing or actuating an environment of the tissue, such as anambient environment and/or an in vivo biological environment. In anembodiment, the photonic device is a tissue-mounted device, for example,a device that is conformally mounted and in physical contact with atissue surface.

Tissue-mounted photonic systems and methods of the invention are capableof generating a range of photonic responses including photonic responsesresulting from an external input, such a photonic response resultingfrom exposure of the device to electromagnetic radiation, for example,as provided by one or more optical sources (e.g., broad band (lamps,LEDs etc.) or narrow band (e.g. a laser)) or ambient light, in opticalcommunication with the device. Photonic responses include opticalresponses corresponding to electromagnetic radiation absorbed, scatteredor emitted by the photonic structures. In an embodiment, for example,the photonic response corresponds to one or more of (i) wavelengths oflight scattered, transmitted or emitted by the photonic structures; (ii)intensity of light scattered, transmitted or emitted by the photonicstructures; (iii) spatial distribution of light scattered, transmittedor emitted by the photonic structures; (iv) phase of light scattered,transmitted or emitted by the photonic structures; and (v) one or morediffraction patterns of light scattered, transmitted or emitted by thephotonic structures. In an embodiment, for example, the photonicresponse corresponds to a measurable change in one or more of: (i)wavelengths of light scattered, transmitted or emitted by the photonicstructures; (ii) intensity of light scattered, transmitted or emitted bythe photonic structures; (iii) spatial distribution of light scattered,transmitted or emitted by the photonic structures; (iv) phase of lightscattered, transmitted or emitted by the photonic structures; and (v)one or more diffraction patterns of light scattered, transmitted oremitted by the photonic structures

A wide range of photonic responses are compatible with the presentphotonic systems. In some embodiments, the photonic response includesspatial and or temporal information corresponding to tissue propertiesand/or environmental properties. Photonic responses of certain systemsof the invention are spatially and/or temporally resolvable responses,for example, reflecting a spatially or temporally varying tissueparameter or environmental parameter. In an embodiment, for example, thephotonic response is a colorimeteric response or fluorometric response,for example, corresponding to the optical characteristics of lightscattered and/or emitted from the photonic structures. In an embodiment,for example, the photonic response is spectroscopic response. In anembodiment, for example, the photonic response results from a change inthe spatial distribution, physical dimensions, phase or chemicalcomposition of the photonic structures. In an embodiment, for example,the photonic response results from a distortion or displacement of thephotonic structures in response to a change in the tissue parameters orenvironmental parameters.

Photonic responses of the present invention are compatible with a rangeof readout modalities including imaging-based optical readout. In anembodiment, for example, a photonic response generated by the presentsystems comprising electromagnetic radiation scattered, absorbed oremitted from the photonic structures is imaged on a camera or otherimaging system, including a CCD, photodiode array or CMOS detector. Inan embodiment, for example, the photonic response is measurable using amobile electronic device, such a photonic response comprisingelectromagnetic radiation scattered, absorbed or emitted from thephotonic structures that is imaged on a camera of a mobile electronicdevice. In some embodiments, for example, the photonic response is adiffraction pattern that is generated by the photonic structures,whereby features of the diffraction pattern correspond to changes intissue parameters or environmental parameters. In an embodiment, asystem of the invention optionally further comprises (i) an opticalsource for illuminating at least a portion of the photonic structuresand/or (ii) an optical detector, such as a camera or other imagingsystem, for detecting electromagnetic radiation scattered, transmittedor emitted from the photonic structures. As used herein, scatteredelectromagnetic radiation is inclusive of scattering at any angleincluding forward and reverse scattering (e.g., reflection). In anembodiment, for example, the photonic response is compatible withcolorimetric, fluorophoric and/or spectroscopic readout, for example,using a mobile electronic device.

In an embodiment, for example, the photonic response corresponds to oneor more tissue parameters selected from the group consisting of: (i)temperature; (ii) hydration state; (iii) chemical composition of thetissue; (iii) chemical composition of a materials derived from thetissue; e.g. a biofluid; (iv) the composition and concentration of ionsof a fluid from the tissue; (iv) pH of a fluid from the tissue; (v) thepresence or absence of a biomarker; (vi) intensity of electromagneticradiation exposed to the tissue; (vii) wavelength of electromagneticradiation exposed to the tissue; and (vii) amount of an environmentalcontaminant exposed to the tissue. In an embodiment, for example, thephotonic response corresponds to one or more environment parametersselected from the group consisting of: (i) intensity of electromagneticradiation exposed to the device; (ii) wavelengths of electromagneticradiation exposed to the device; (iii) amount of an environmentalcomponent exposed to the device; (iv) chemical composition of anenvironmental component exposed to the device; (v) amount of anenvironmental contaminant exposed to the device; (vi) chemicalcomposition of an environmental contaminant exposed to the device. In anembodiment, the photonic response is an optical signal.

A wide range of photonic structures are useful in the present systemsand methods. In an embodiment, for example, the one or more photonicstructures optically absorb, scatter, transmit or emit electromagneticradiation having wavelengths in the visible, ultraviolet or infraredregions of the electromagnetic spectrum. In an embodiment, use ofvisible region (e.g. 350 nm to 750 nm) and near-IR region (e.g.,750-1300 nm) of the electromagnetic spectrum light is preferred tominimize an potential adverse effects to the tissue. In an embodiment,the electromagnetic radiation exposed to the photonic device and/theelectromagnetic radiation scatter or emitted from the photonic device ischaracterized by wavelengths selected over the range of 350 nanometersto 1300 nanometers, and optionally wavelengths selected over the rangeof 400 nanometers to 900 nanometers.

In an embodiment, for example, the one or more photonic structures areflexible or stretchable photonic structures, for example, exhibitingstretchability, without mechanical failure and/or degradation of opticalproperties, of greater than or equal to 5%, and greater than or equal50% for some embodiments and greater than or equal 100% for someembodiments. In an embodiment, for example, the one or more photonicstructures are microstructures (e.g., having physical dimensionsselected from the range of 1 micron to 1000 microns) and/ornanostructures (e.g., having physical dimensions selected from the rangeof 1 nm to 1000 nm). In an embodiment, for example, the one or morephotonic structures are characterized by an average modulus less than orequal to 100 MPa, optionally for some embodiments less than or equal 500kPa. In an embodiment, for example, the one or more photonic structuresare characterized by an average modulus selected over the range of 0.5kPa to 100 MPa, optionally for some applications selected over the rangeof 0.5 kPa to 500 kPa. In an embodiment, for example, the one or morephotonic structures are characterized by average lateral dimensionsselected from the range of 10 μm to 1 cm and/or average thicknessselected from the range of 1 μm to 1000 μm, optionally for someembodiments, average lateral dimensions selected from the range of 10 μmto 1000 μm and/or average thickness selected from the range of 1 μm to100 μm. In an embodiment, for example, the one or more photonicstructures are capable of mechanical deformation in response to astimulus, such as a change in temperature. In an embodiment, forexample, at least a portion of the one or more photonic structures arein fluid communication, thermal communication, optical communication,and/or electrical communication with the tissue. In an embodiment, forexample, at least a portion of the one or more photonic structures arein physical contact with the surface of the tissue.

Useful photonic structures for some embodiments of the present systemsand methods are spatially distributed in an array, such as an array withindividual photonic structures individually in physical, optical orthermal contact with specific regions of the tissue surface. Photonicstructures provided in an array form factor is useful in certain systemsand methods to provide a photonic response characterizing spatialinformation corresponding to the tissue or environment, such as aspatial distribution of tissue parameters or environmental parameterswith respect to a tissue surface. In an embodiment, for example, thearray of photonic structures is a pixelated array; wherein each photonicstructure independently corresponding to an individual position thearray. In an embodiment, for example, the array of photonic structuresis a pixelated array, for example positions in the array individuallyaddressed to specific regions of the tissue surface.

In an embodiment, for example, individual pixels or the array have anaverage lateral dimensions selected from the range of 10 μm to 1000 μm,optionally for some embodiments selected from the range of 100 μm to 500μm and further optionally for some embodiments selected from the rangeof 200 μm to 500 μm. In an embodiment, for example, the individualpixels have an average thickness selected from the range of 1 μm to 100μm, optionally for some embodiments selected from the range of 10 μm to100 μm and further optionally for some embodiments selected from therange of 20 μm to 50 μm. In an embodiment, for example, the individualpixels are spaced from adjacent pixels in the array other by a distanceselected from the range of 10 μm to 1000 μm, optionally for someembodiments a distance selected from the range of 100 μm to 1000 μm andfurther optionally for some embodiments a distance selected from therange of 250 μm to 500 μm. In an embodiment, for example, the pixelatedarray comprises 10 to 1,000,000 pixels, optionally for some embodiments10 to 100,000 pixels. In an embodiment, for example, the pixelated arrayhas a footprint selected from the range of 10 mm² to 2000 cm².

Photonic structures useful in the present systems and methods includestructures incorporating optical indicators, such as colorimetric orfluorometric indicators, having optical properties that are useful forcharacterizing tissue parameters or environmental parameters. In anembodiment, for example, at least a portion of the pixels comprise acolorimetric indicator, fluorometric indicator or both, including deviceincluding pixels corresponding to different colorimetric and/orfluorometric indicators. The invention is compatible with a range ofphotonic structures incorporating indicators including embedded and/orencapsulated structures. In an embodiment, for example, the photonicstructures are microencapsulated structures and/or nano-encapsulatedstructures, for example, having an indicator that is encapsulated by oneor more encapsulation structures, such as laminating, embedding orencapsulation layers. In an embodiment, the microencapsulated structuresand/or nano-encapsulated structures are in physical, thermal, optical orelectrical contact with the tissue of a material(s) derived from thetissue, such as a biofluid.

In an embodiment, for example, at least a portion of the pixels comprisea colorimetric indicator that is a liquid crystal, an ionochromic dye, apH indicator, a chelating agent, a fluorphore or a photosensitive dye.In an embodiment, for example, at least a portion of the pixels comprisea colorimetric indicator capable of generating a photonic response forcharacterizing a temperature, exposure to electromagnetic radiation or achemical composition of a tissue or material derived from tissue. In anembodiment, for example, at least a portion of the pixels comprise acolorimetric indicator comprising a thermochromic liquid crystal thatunder goes a measurable change in the wavelength of light that isabsorbed, transmitted or scattered upon a change of the tissueparameter. In an embodiment, for example, at least a portion of thepixels comprise a colorimetric indicator comprising chiral nematicliquid crystal that undergoes a measurable change in the wavelength oflight that is absorbed, transmitted or scattered upon a change intemperature of the tissue.

In an embodiment, for example, at least a portion of the pixels comprisea colorimetric indicator comprising an ionochromic dye that under goes ameasurable change in the wavelength of light that is absorbed,transmitted or scattered in response to a composition or property of thetissue or a material derived from the tissue such as a biological fluid.In an embodiment, for example, the composition or property of thebiological fluid corresponds to a change in pH, concentration of freecopper ion, or concentration of iron ion. In an embodiment, for example,at least a portion of the pixels comprise a colorimetric indicator thatundergoes a measurable change in color in response to exposure toultraviolet radiation. In an embodiment, for example, the photonicstructures include colorimetric or fluorometric indicators that changeoptical properties upon contact with a biomarker in the tissue or in amaterial derived from the tissue such as a biological fluid

In an embodiment, for example, the pixelated array further comprises oneor more calibration pixels, such as dots having a fixed color.

A range of stretchable and flexible substrates are useful in embodimentsof the present photonic devices and methods. In some embodiment, thesubstrate is a functional substrate. Use of low modulus and thinsubstrates are beneficial in some embodiments for achieving a conformalcontact with tissue surface having complex morphologies withoutdelamination and achieving a conformal contact without movement of thedevice relative to the contact surface of the tissue, for example,during movement of tissue. Use of selectively colored or opticallyopaque substrates are useful for providing contrast sufficient foreffective optical readout, for example, via imaging using a mobileelectronic device. Use of porous substrates and substrates havingfluidic structures (e.g., active or passive fluidic channels) arebeneficial for embodiments capable of characterizing properties offluids from the tissue.

In an embodiment, for example, the substrate is optically opaque. In anembodiment, for example, the flexible or stretchable substrateincorporates one or more fluidic structures for collecting ortransporting fluid from the tissue to the photonic structures. In anembodiment, for example, the flexible or stretchable substrate comprisesan elastomer. In an embodiment, for example, the flexible or stretchablesubstrate is a low modulus rubber material or a low modulus siliconematerial. In an embodiment, for example, the flexible or stretchablesubstrate is a bioinert or biocompatible material. In an embodiment, forexample, the flexible or stretchable substrate comprises a gas-permeableelastomeric sheet. In an embodiment, for example, the flexible orstretchable substrate has an average modulus less than or equal to 100MPa, optionally for some embodiments less than or equal to 500 kPa,optionally for some embodiments less than or equal to 100 kPa. In anembodiment, for example, the flexible or stretchable substrate has anaverage modulus selected over the range of 0.5 kPa to 100 MPa, andoptionally for some embodiments 0.5 kPa to 500 kPa, and optionally forsome embodiments 0.5 kPa to 100 kPa. In an embodiment, for example, theflexible or stretchable substrate has an average thickness less than orequal to 3 mm, and for some applications less than or equal to 1000microns. In an embodiment, for example, the flexible or stretchablesubstrate has an average thickness selected over the range of 1 to 3000microns, and for some applications 1 to 1000 microns.

Photonic devices of the invention may further comprise a range ofadditional device components. In an embodiment, for example, the devicefurther comprises one or more additional device components supported bythe flexible or stretchable substrate, the device components selectedfrom the group consisting of an electrode, strain gauge, optical source,temperature sensor, wireless power coil, solar cell, wirelesscommunication component, photodiode, microfluidic component, inductivecoil, high frequency inductor, high frequency capacitor, high frequencyoscillator, high frequency antennae, multiplex circuits,electrocardiography sensors, electromyography sensors,electroencephalography sensors, electrophysiological sensors,thermistors, transistors, diodes, resistors, capacitive sensors, andlight emitting diodes. In an embodiment, for example, the device furthercomprises one or more wireless communication antenna structures ornear-field communication coils supported by the flexible or stretchablesubstrate. In an embodiment, for example, the device further comprisesone or more single crystalline semiconductor structures supported by theflexible or stretchable substrate.

In an embodiment, for example, the device further comprises one or moreoptical components supported by the stretchable or flexible substrate,and optionally providing in optical communication of the photonicstructures. In an embodiment, for example, the optical components areone or more of a light collecting optical component, a lightconcentrating optical component, a light diffusing optical component, alight dispersing optical component and a light filtering opticalcomponent. In an embodiment, for example, the optical components are oneor more of a lens, a lens array, a reflector, an array of reflectors, awaveguide, an array of waveguides, an optical coating, an array ofoptical coatings, an optical filter, an array of optical filters, afiber optic element and an array of fiber optic elements.

In some embodiment, the photonic structures are in physical contact withthe substrate. Photonic devices of the invention include multilayerdevices, for example, including one or more additional layer such asencapsulating layers at least partially encapsulating the photonicstructures, and/or intermediate layers provided between the one or morephotonic structures and the substrate. In an embodiment, the photonicstructures are provided proximate to a neutral mechanical surface of thedevice. In an embodiment, for example, the photonic structures arepositioned proximate to a neutral mechanical surface of the device, suchas provided distance less than 2 mm, less than 10 μm, less than 1 μm, orless than 100 nm to a neutral mechanical surface. In an embodiment, forexample, the thickness and/or physical properties (e.g., Young'smodulus) of substrate and encapsulating layers are selected to positionthe photonic structure positioned proximate to a neutral mechanicalsurface of the device.

The device level mechanical, thermal, electronic and optical propertiesof the present photonic devices is important for supporting a range oftechnology applications. In an embodiment, for example, the device has amodulus within a factor of 1000, and optionally a factor of 10, of amodulus of the tissue at the interface with the device. In anembodiment, for example, the device has an average modulus less than orequal to 100 MPa, optionally for some embodiments less than or equal to500 kPa, optionally for some embodiments less than or equal to 200 kPaand optionally for some embodiments less than or equal to 100 kPa. In anembodiment, for example, the device has an average modulus selected overthe range of 0.5 kPa to 100 MPa, optionally for some embodimentsselected over the range of 0.5 kPa to 500 kPa, optionally for someembodiments selected over the range of 1 kPa to 200 kPa.

Matching the physical dimensions and properties of the devices to thatof the tissue is a useful design strategy in some embodiments to achieverobust conformal contact. In an embodiment, for example, the device hasan average modulus equal to or less than 100 times, optionally equal toor less than 10 times, the average modulus of the tissue at theinterface. In an embodiment, for example, the device has an averagethickness less than or equal to 3000 microns, optionally for someembodiments less than or equal to 1000 microns. In an embodiment, forexample, the device has an average thickness selected over the range of1 to 1000 microns. In an embodiment, for example, the device has a netbending stiffness less than or equal to 1 mN m, optionally for someembodiments less than or equal to 1 nN m, optionally for someembodiments less than or equal to 0.1 nN m and optionally for someembodiments less than or equal to 0.05 nN m. In an embodiment, forexample, the device has a net bending stiffness selected over the rangeof 0.01 nN m to 1 N m, optionally for some applications selected overthe range of 0.01 to 1 nN m, and optionally for some embodimentsselected over the range of 0.1 to 1 nN m. In an embodiment, for example,the device has an areal mass density less than or equal to 100 mg cm⁻²,optionally for some applications less than or equal to 10 mg cm⁻². In anembodiment, for example, the device has an areal mass density selectedover the range of 0.1 mg cm⁻² to 100 mg cm⁻², optionally for someapplications elected over the range of 0.5 mg cm⁻² to 10 mg cm⁻². In anembodiment, the device is characterized by a stretchability greater thanor equal to 5% and optionally for some applications 50% and optionallyfor some applications 100%, for example, by being able to undergostretching to this extent without mechanical failure. In an embodiment,the device is characterized by a stretchability selected from the rangeof 5% to 200% and optionally for some applications 20% to 200%, forexample, by being able to undergo stretching to this extent withoutmechanical failure.

The photonic systems of the invention are compatible with a range oftissue types including in vivo tissues, internal tissues and externaltissues. In some embodiments, the tissue is skin, heart tissue, braintissue, muscle tissue, nervous system tissue, vascular tissue,epithelial tissue, retina tissue, ear drum, tumor tissue, or digestivesystem structures. In some embodiments, for example, the deviceestablishes conformal contact with the tissue when the device is placedin physical contact with the tissue, and wherein the conformal contactwith the tissue in the biological environment is maintained as thetissue moves or when the device moves. The tissue may be of a subjectthat is undergoing treatment or diagnosis. In some embodiments, forexample, the device is capable of establishing conformal contact withthe tissue surface in the presence of a biofluid.

In an aspect, the invention provides a method of sensing one or moretissue parameters or environmental parameters, the method comprising thesteps of: (i) providing the tissue of the subject; (ii) contacting asurface of the tissue with a photonic device, wherein the photonicdevice comprises: (1) a flexible or stretchable substrate; and (2) oneor more photonic structures supported by the flexible or stretchablesubstrate for generating a photonic response corresponding to said oneor more tissue parameters or environmental parameters; wherein theflexible or stretchable substrate and the one or more photonicstructures provide a net bending stiffness (and/or Young's modulus) suchthat the device establishes conformal contact with a surface of thetissue; and (3) detecting the photonic response from the photonicdevice, thereby sensing the one or more tissue parameters orenvironmental parameters. Methods of this aspect may further includedetecting the photonic response using a two-dimensional optical detectorcapable of spatially resolving the photonic response, such as a cameraor other imaging device including using a mobile electronic device.Methods of this aspect may further include detecting the photonicresponse as a function of time. In an embodiment, for example, the stepof measuring the photonic response from the photonic device comprisesdetecting electromagnetic radiation scattered or emitted by the one ormore photonic structures. In an embodiment, for example, detectingelectromagnetic radiation scattered or emitted by the one or morephotonic structures is carried out using a mobile electronic device.Methods of this aspect may further comprise generating a detector signalcorresponding to the photonic response using said optical detector.Methods of this aspect may further comprise analyzing the detectorsignal, thereby determining said one or more tissue parameters orenvironmental parameters.

Embodiments of this aspect include the step of establishing conformalcontact with one or more surfaces of the tissue. In an embodiment, forexample, the photonic device is provided in in conformal contact withtissue selected from the group consisting of: skin, heart tissue, braintissue, muscle tissue, nervous system tissue, vascular tissue,epithelial tissue, retina tissue, ear drum, tumor tissue, and digestivesystem structures. In an embodiment, for example, the tissue is skin andwherein the device establishes conformal contact with the outer surfaceof the epidermis. The methods of the invention include the step ofcontacting tissue of a subject with the photonic device, such as a humansubject or other animal. In some embodiments, subjects of the presentmethods refer to a subject (1) having a condition able to be monitored,diagnosed, prevented and/or treated by administration of photonic deviceof the invention; or (2) that is susceptible to a condition that is ableto be monitored, diagnosed, prevented and/or treated by administering aphotonic device of the invention.

Without wishing to be bound by any particular theory, there may bediscussion herein of beliefs or understandings of underlying principlesrelating to the devices and methods disclosed herein. It is recognizedthat regardless of the ultimate correctness of any mechanisticexplanation or hypothesis, an embodiment of the invention cannonetheless be operative and useful.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A shows an embodiment of a photonic device for interfacing with atissue in a biological environment, including for example a tissuemounted device. FIG. 1B shows an embodiment of a method of sensing oneor more tissue parameters of a tissue of the subject or environmentalparameters.

FIG. 2. Pictures, micrographs and design features of an ‘epidermal’thermochromic liquid crystal (e-TLC) thermal imaging device. a, Pictureof devices deformed by pinching the skin in a twisting motion (left),poking with a warm glass rod while on skin (middle) and collapsing underits own weight while free-standing (right). b, Magnified view of adevice operating in the blue region of the spectrum, without (top) andwith (bottom) integrated patterns of dots that have fixed colors forcalibration. c, Picture of an e-TLC device with calibration system,operating the curved surface of the skin. d, Picture of a device thatincludes a radio frequency antenna and Joule heating element on its backsurface, folded over and resting on palm, with an enlarged view of theserpentine antenna structure (inset). e, Schematic illustration offinite element modeling results for an e-TLC device with wireless heaterunder tensile strain, with magnified view of the Joule heating element(inset). f, Image of an active, wireless e-TLC device collected whileexposed to RF power in air, with magnified view of the color changesinduced by the heater (inset). g, Infrared image of the same deviceunder similar conditions, with magnified view in the region of theheater (inset).

FIG. 3. Experimental and computational studies of the mechanicalproperties of e-TLC devices. a, Measurements and theoreticalcalculations of stress-strain responses of a device. b, Comparisonbetween images and three dimensional finite element modeling of arepresentative region of e-TLC device under different levels of tensilestrain.

FIG. 4. Calibration and use of e-TLC devices for precision thermalimaging on the skin. a, Reflectance measured at a single pixel from 32°C. to 39° C. and corresponding images for 33° C. to 38° C. (inset). b,Temporal variations in temperature extracted from digital color analysisof an e-TLC held, nominally, at a constant temperature. c, Temporalvariations in apparent temperature determined from color analysis ofcalibration pixels in an e-TLC device. Frames b and c also show resultsobtained with an infrared camera. d, Illustration of the steps forprocessing digital images of e-TLC devices, demonstrated on arepresentative 7×7 array of pixels. e, Color-temperature calibrationdetermined using hue analysis. f, Images of a e-TLC device that consistsof an 26×26 array of pixels, conformally mounted on the wrist. g, 3Drendering of the temperature distribution extracted from the colorinformation obtained by hue value analysis of digital images of thedevice. h, 2D rendering of temperature captured by an infrared camera atthe same time and at the same location as in g. i, Line-cut temperatureprofiles extracted from the data of g and h.

FIG. 5. Temperature analysis with an e-TLC device that incorporates anarray of color calibration pixels co-located with sensing pixels,evaluated under different illumination conditions. a, Image of a deviceimmediately after local heating at the center of the array. b, Hue andsaturation values extracted for the calibration (stars) and sensingpixels (dots; red—illumination with a fluorescent light;blue—illumination with a light emitting diode; green—illumination with ahalogen lamp). 3D rendering of color-corrected temperatures determinedwith c, white fluorescent light (FL), d, white light-emitting diode(LED), e, halogen light (HG). f, Line graphs of results collected alongthe dashed lines shown in c-e. g, Results similar to those in f, butwithout color correction.

FIG. 6. Determination of thermal conductivity and thermal diffusivity ofthe skin using active e-TLC devices. a, Example of temperatures(symbols) as a function of distance from the position of local heatingin an active e-TLC device and corresponding best fit modeling results(analytic; line), for determining the thermal conductivity. b, Thermalconductivity of water/ethylene glycol solutions evaluated using anactive e-TLC device, with comparison to values obtained from theliterature and from analysis of temperatures determined with an infraredcamera. c, Thermal conductivities measured with an active e-TLC deviceon the skin at different levels of hydration, separately measured with acommercial moisture meter. The error bars represent average standarddeviations of measurements obtained with the moisture meter. d, Exampleof temperatures (symbols) as a function of time for a location near awireless heater in an active e-TLC device, and corresponding best fitmodeling results (analytic; line) for determining the thermaldiffusivity. e, Thermal diffusivity of water/ethylene glycol solutionsevaluated using an active e-TLC device, with comparison to valuesobtained from the literature and from analysis of temperaturesdetermined with an infrared camera. f, Thermal diffusivities measuredwith an active, wireless e-TLC device on the skin at different levels ofhydration, separately measured with a commercial moisture meter. Theerror bars represent average standard deviations of measurementsobtained with the moisture meter.

FIGS. 7A-7G. Application of an e-TLC thermal imaging device in areactive hyperaemia test. FIG. 7A shows optical images of an e-TLCdevice on the wrist during an occlusion test after blood is released(left) with magnified view (right). FIG. 7B shows infrared image of thedevice (left) with magnified view (right). FIG. 7C shows 3D rendering ofspatial distributions of temperature determined with the e-TLC device atdifferent times during and after occlusion (occlusion starts at t=0 sand ends at t=160 s). FIG. 7D shows line graphs of temperatures alongthe horizontal dashed red line in the right frame of FIG. 7A, at varioustimes. FIG. 7E shows line graphs of temperatures along the verticaldashed red line in the right frame of FIG. 7A, at various times. FIG. 7Fshows rate of blood flow through the ulnar artery determined bycomparison of thermal models to experimental results. The key parametersinclude: the occlusion time (tocc)=160 s; time-to-peak-flow (tdw)=15 s;the baseline flow rate (ω0)=30 mL/min; the occluded flow rate (ωs)=1.5mL/min; and the peak flow rate (ωmax)=90 mL/min. FIG. 7G shows measuredtemperature rise at the surface of the skin above the ulnar arteryduring the occlusion along with results from finite element analyses(FEA) using the blood flow rate in FIG. 7F. The eight sub-framescorrespond to the temperature histories of different points at thehorizontal dashed red line in the right frame of FIG. 7A.

FIG. 8. Process for fabricating e-TLC devices. (a) A PDMS stamp with anarray of posts embossed on its surface is ‘inked’ by bringing it intocontact with a uniform layer of TLC aqueous slurry spin cast on a glassslide while still in wet state. The thickness of the ink was ˜100 μm toensure that the ink contacts on the top surfaces of the posts. (b) Theinked TLC material on the PDMS stamp was allowed to dry in air for 15minutes. The thickness of the dried film is ˜15 μm. Additional ‘inking’processes are repeated to achieve a final thickness of 25-30 μm. Atypical TLC pixel is thickest in the center due to the hydrophobicnature of the PDMS surface and the large contact angle formed during theinking process. (c) Transfer printing allows delivery of the TLC to apiece of thermal release tape. (d) Transfer to the black PDMS substrateis enabled by heat activated release from the tape. (e) The device isencapsulated with a transparent layer of PDMS by spin casting.

FIG. 9. Mechanical response of an e-TLC device to uniaxial strain. (a)Experimental, analytical and finite element modeling results for thechange in horizontal and vertical spacings between adjacent pixels underdifferent levels of tensile strain. (b) Comparison between images andthree dimensional finite element modeling of a representative region ofan e-TLC device that incorporates color calibration pixels underdifferent levels of tensile strain.

FIG. 10. Experimental and computational studies of the mechanicalproperties of Joule heater element. (a) Comparison between experimentalimages and three dimensional finite element modeling of a wired Jouleheating element under different levels of tensile strain, and straindistribution computed for the case of stretching to 50%. (b) Comparisonbetween experimental images and three dimensional finite elementmodeling of a wireless Joule heater under different levels of tensilestrain, and strain distribution computed for the case of stretching to50%.

FIG. 11. Experimental studies of the effect of strain on the efficiencyof wireless Joule heating. (a) Infrared temperature measurements for awireless Joule heater under exposure to RF energy while mechanicallydeformed in different ways, both in air and on skin. (b) Measurements atdifferent levels of tensile strain with corresponding images.

FIG. 12. Water permeability test. (a) Images of the experimental set-upsfor measurement of water permeation according to ASTM E96-95 guidelines,and (b) Results of the change in weight as a function of time associatedwith water uptake by the dessicant, for e-TLC devices with differentthicknesses and for a commercial TLC strip.

FIG. 13. Effect of e-TLC operation on temperature and hydration of theskin. (a) Infrared image captured immediately after mounting an e-TLCdevice on the wrist. (b) Infrared image captured 3 hours after mounting.For both (a) and (b), the data indicate that the average temperatures atthe regions of the device are the same as those adjacent to the device.(c) Temperature difference between a point near the device and a pointunderneath the device shows no obvious increase during the three houroperation. (d) Hydration level read from a commercial hydration metershows a maximum increase of about 25% after 3 hours operation on a verydry skin.

FIG. 14. Sensor response time. (a) Layers used in analytical modeling todetermine sensor response time on skin. (b) Experimental setup formeasuring sensor response time. A warm ethylene glycol bath, which hassimilar thermal properties to skin, is in contact with the e-TLC devicefrom the back surface. (c) Experimental sensor response time captured byhigh speed camera, and corresponding analytic predictions based on aone-dimensional heat conduction model. In experiment, the time requiredfor the sensor to reach 90% of the total temperature change is achievedin one frame which is approximate 33 ms for the case of 30 μm black PDMSand 25 μm liquid crystal.

FIG. 15. Noise and uncertainty examined using temperature insensitiveacrylic colors. (a) TLC color-temperature calibration plotted in thehue/saturation space. Symbols are located at positions corresponding tothe hue/saturation values of the TLC during calibration runs, asindicated with their hue values. Temperatures are calculated with a twodimensional linear fit and are represented by a color gradient. (b)Temporal fluctuation in the color of the TLC, when held at a nominallyfixed temperature. (c) Temporal fluctuation of the blue calibrationcolor at fixed temperature. (d) Temporal fluctuation of the greencalibration color at fixed temperature. (e) Temporal fluctuation of thered calibration color at fixed temperature.

FIG. 16. Finite element models that allow determination thermalconductivity and diffusivity from data collected using active e-TLCdevices. (a) A 3D view of a model with a Joule heater embedded betweenan e-TLC device and the skin. (b) A cross-sectional view of a model witha Joule heater embedded between an e-TLC device and the skin. (c)Analytical model of the spatial decay in temperature at steady stateduring operation of the Joule heater. (d) Corresponding finite elementmodeling results. (e) Analytical and finite element model of the spatialtemperature decay with a wired Joule heater operation along onedimension. (f) Analytical and finite element model of the temporaltemperature rise with a wireless Joule heater operation for locationsaway from the heater. (g) Skin heat capacity inferred from the skinthermal conductivity and diffusivity values in FIG. 6.

FIG. 17. Comparison of an e-TLC thermal imaging device and infraredcamera measurement in a reactive hyperaemia test. (a) Spatialdistributions of temperature determined with the e-TLC device atrepresentative times from t=160 s to t=260 s at an interval of 20 s. (b)Spatial distributions of temperature determined with the infrared cameraat representative times from t=160 s to t=260 s at an interval of 20 s.

FIG. 18. Schematic illustration of the thermal conduction model thatdetermines the blood flow rate during occlusion. (a) Cross-sectionalview and (b) three-dimensional view of the wrist model; (c) Temperaturevariance of FEA and experiment versus the baseline flow rate; (d)Experimental results of the steady-state temperature as a function ofthe distance from the artery, as compared to the FEA calculations usingthe baseline flow rate of 30 mL/min; (e) Distribution of temperaturevariance in the space of parameters, a and τ₀, during stage II ofocclusion.

FIG. 19. (a) Schematic illustration of a passive wireless capacitivesensor designed for sensing of sweat from the surface of the skin.Pictures of a device in (b) longitudinal and (c) latitudinal states ofdeformation, and crumpled between the fingers (d). Pictures of a devicemounted on the skin in (e) undeformed, (f) uniaxially stretched and (g)biaxially stretched configurations.

FIG. 20. (a) Scanning electron micrograph of a sensor on a PUR substratecoated with a thin silicone film; the regions colorized in yellowrepresent the interdigitated gold electrodes. (b) Picture of a sweatsensor and a reference sensor on the arm of a volunteer for in-vivotesting. (c) Picture of a sweat sensor underneath a primary coil. Asyringe needle inserted into the sensor delivers controlled amounts of abuffer solution through a syringe pump. (d) Representative data showingthe response of the sensor (resonant frequency, f₀) as a function oftime after introduction of 0.6 mL buffer solution (labeled 1). Theinitial response (labeled 2) corresponds to wicking of the solution intothe porous substrate, to yield a stable overall shift in f₀ (labeled 3).As the solution evaporates over the next several hours, f₀ recovers toapproximately the initial value. The inset shows the phase differencemeasured by the primary coil at the three time points indicated in themain frame. (e, f) Results of testing on two volunteers, withcomparisons to changes in weight evaluated using similar poroussubstrates (without detection coils) placed next to the sensors. Both f₀and the weight of the sensors calibrated from f₀ are shown, along withcomparison to the weight of the reference substrates. (g) Phase responseof a sensor under biaxial strain from 0 to 27%. (h) Phase response as afunction of concentration of sodium chloride, from 0 to 4.5 g/L. (i)Change in f₀ of a sweat sensor on a CP substrate as a function of timeduring controlled injection of 0.6 mL buffer solution.

FIG. 21. (a) Wireless sweat sensors based on different poroussubstrates. (b) SEM images of the substrates coated with thin layer ofsilicone to facilitate chemical bonding between the sensors and thesubstrates. (c) Weight gain of different substrate materials associatedwith immersion in water. (d) Porosity of the substrate materials. (e)Images of strips of the substrate materials when partially immersed intowater with red dye. (f) Water permeability of the substrate materials.

FIG. 22. (a) Images that illustrate a simple colorimetric detectionscheme, based on systematic increases in transparency with waterabsorption. (b) The ratio of RGB intensity for a sensor like the oneillustrated in (a), as a function of water absorption. (c) An image andvector diagrams corresponding to a sensor and its expansion due to waterabsorption. (d) Series of pictures of a sensor doped with a pHindicator, each collected with absorbed water at a different pH value.(e) Absorbance of RGB channels at different pH values. (f) Absorbance ofRGB channels at different copper concentrations. (g) Absorbance of RGBchannels at different iron concentrations.

FIG. 23. (a) Capacitance values of a coaxial cable probe when in contactwith sensors on CP and PUR substrates injected with 0.6 mL buffersolution. (b) Stability of a sweat sensor at temperatures from 25 to 45°C. (c) Time variation of f₀ for a sweat sensor on a silicone substratein response to the injection of 0.6 mL buffer solution. (d) Drift andstability of a sensor output at dry state over an extended period of 3hours.

FIG. 24. (a) A sensor is biaxially stretched by two perpendicularstretchers at a strain from 0 to 27%. (b) Expansion of the surface areaof the sensor in response to water absorption.

FIG. 25. (a) SEM images of porous materials, showing that the pores ofPUR and Silicone dressing are uniform and that the pores of RCS, PVAS,and CP are amorphous. (b) Contact angle measurements performed bypartially immersing strips of the porous materials into water dyed withred color, and recording the angle at the interface of two materials.

FIG. 26. (a) Color changes in the sensor when the free copperconcentration changes from 0 to 1 mg/L, (b) Color changes in the sensorwhen the iron concentration changes from 0 to 0.8 mg/L.

FIG. 27. (a)-(g) Fabrication processes for a wireless sweat sensor.

FIG. 28. Exploded view of a colorimetric sensor comprising a near-fieldcommunication coil.

FIG. 29. Photograph of the device of FIG. 46 adhered to the skin of asubject.

FIG. 30. Fabrication method and adhesion test on skin.

FIG. 31. Artificial sweat pore test using a syringe to feed artificialsweat at a rate of 12 μL/hr.

FIG. 32. Colorimetric detection of various biomarkers using a sweatsensor for self-monitoring and early diagnosis.

FIG. 33. Absorbance spectrum illustrating the color change of a reactantthat may be used to determine sweat volume and rate.

FIG. 34. Absorbance spectrum and legend illustrating the color change ofa reactant(s) that may be used to determine sweat pH, which may becorrelated with sodium concentration, indicating to a user the propertime to hydrate.

FIG. 35. Absorbance spectrum and legend illustrating the color change ofa reactant(s) that may be used to determine glucose concentration insweat, which may be correlated with blood glucose concentration.

FIG. 36. Absorbance spectrum and legend illustrating the color change ofa reactant(s) that may be used to determine lactate concentration insweat, which may provide an indication of shock, hypoxia and/or exerciseintolerance.

FIG. 37. A sweat sensor incorporating colorimetric biomarker indicatorsprovides qualitative and quantitative data that may be observed by thenaked eye and/or wirelessly observed by a detection device, such as asmartphone.

FIG. 38. (A) Schematic illustration of an epidermal microfluidic sweatsensor providing information of sweat volume and rate as well asconcentration of biomarkers in sweat incorporated with wirelesscommunication electronics. (B) Fabrication process for flexible andstretchable epidermal microfluidics. (C) Pictures of fabricated sweatsensors mounted on the skin under various mechanical stresses.

FIG. 39. (A) Picture of fabricated epidermal sweat sensor indicatinginformative detection schemes for sweat analysis. (B) In vitroartificial sweat pore system set up. (C) Optical image of sweat sensorapplied on artificial pore membrane. (D) Scanning electron microscopy(SEM) image of the artificial pore membrane. Inset shows magnified imageof single pore. (E) Representative images of sweat patch on theartificial sweat pore system while mimicking sweating events for 5 h.Sweat flowed continuously in the microfluidic systems along with colorchange accordingly.

FIG. 40. Analytical colorimetric detections and respective UV-Visspectrums of biomarkers in sweat. (A) Spectrum of anhydrous (blue) andhexahydrate (pale pink) cobalt (II) chloride. The presented color in thespectrum corresponds to the observed color with naked eye. (B) Opticalimages of resulted color change of the filter papers as a function ofvarious pH values and analyte concentrations. (C) Spectrum of universalpH assay with various buffer solutions in the range of pH 5.0-8.5. (D-F)Spectrum of biomarkers in sweat as a function of concentration ofanalytes: glucose (D), lactate (E) and chloride (F). The presented colorfor each spectrum corresponds to exhibited color on paper-basedcolorimetric results, which is presented in image (B). Insets indicatecalibration curves of respective analytes corresponding withconcentration in the optical images (B). All spectra were determined atroom temperature.

FIG. 41. (A) An image of fabricated sweat sensor incorporated withnear-field communication electronics. (B) Demonstration pictures ofwireless communication via smartphone. The RGB information wasdetermined using an android image analysis app.

FIG. 42. (A) Schematic illustration of an epidermal microfluidic sweatsensor providing information on sweat volume and rate as well asconcentration of biomarkers in sweat incorporated with wirelesscommunication electronics and an adhesive layer. (B) Schematicillustration of image process markers applied to an epidermalmicrofluidic sweat sensor.

FIG. 43. Graphical representation of water loss as a function of outletchannel (A) width and (B) length.

FIG. 44. Graphical representation of back pressure inside a channelshowing that shorter outlet channels and larger channel widths producelower back pressures.

FIG. 45. (A) Schematic illustration of a cross section of a microfluidicchannel deformed due to pressure. (B) Schematic illustration of a topperspective view of a section of an epidermal microfluidic sweat sensorshowing a width of the microfluidic channel. (C) Graphicalrepresentation of deformation shown as volume change due to pressure.

FIG. 46. (A) Experimental set-up for 90° peel adhesion property testing(standard ISO 29862:2007) using a force gauge (Mark-10, Copiague, N.Y.).Images of (B) holding devices adhered on the skin with a force gauge and(C) peeling devices at an angle of 90°. (D) Force measurement whiledisplacing the device at a rate of 300 mm/min indicated by the grayregion where peeling occurs. Determined average peeling force is 5.7 N.

FIG. 47. Colorimetric determination of creatinine. (A) UV-VIS spectrumwith various creatinine concentrations (i.e., 15-1000 μM) and (B)constructed calibration based on this spectrum. The presented color foreach spectrum corresponds to exhibited color on paper-based colorimetricdetection reservoirs as a function of creatinine concentration, which ispresented in optical image (C).

FIG. 48. Colorimetric determination of ethanol. (A) UV-VIS spectrum withvarious ethanol concentrations (i.e., 0.04-7.89% (w/v)) and (B)constructed calibration based on this spectrum. The presented color foreach spectrum corresponds to exhibited color on paper-based colorimetricdetection reservoirs as a function of ethanol concentration, which ispresented in optical image (C).

FIG. 49. Various microfluidic sweat sensor designs.

FIG. 50. Various types of orbicular channel designs and respectivelycalculated channel properties.

DETAILED DESCRIPTION OF THE INVENTION

In general, the terms and phrases used herein have their art-recognizedmeaning, which can be found by reference to standard texts, journalreferences and contexts known to those skilled in the art. The followingdefinitions are provided to clarify their specific use in the context ofthe invention.

“Functional substrate” refers to a substrate component for a devicehaving at least one function or purpose other than providing mechanicalsupport for a component(s) disposed on or within the substrate. In anembodiment, a functional substrate has at least one skin-relatedfunction or purpose. In an embodiment, a functional substrate of thepresent devices and methods exhibits a microfluidic functionality, suchas providing transport of a bodily fluid through or within thesubstrate, for example via spontaneous capillary action or via an activeactuation modality (e.g. pump, etc.). In an embodiment, a functionalsubstrate has a mechanical functionality, for example, providingphysical and mechanical properties for establishing conformal contact atthe interface with a tissue, such as skin. In an embodiment, afunctional substrate has a thermal functionality, for example, providinga thermal loading or mass small enough so as to avoid interference withmeasurement and/or characterization of a physiological parameter, suchas the composition and amount of a biological fluid. In an embodiment, afunctional substrate of the present devices and method is biocompatibleand/or bioinert. In an embodiment, a functional substrate may facilitatemechanical, thermal, chemical and/or electrical matching of thefunctional substrate and the skin of a subject such that the mechanical,thermal, chemical and/or electrical properties of the functionalsubstrate and the skin are within 20%, or 15%, or 10%, or 5% of oneanother.

In some embodiments, a functional substrate that is mechanically matchedto a tissue, such as skin, provides a conformable interface, forexample, useful for establishing conformal contact with the surface ofthe tissue. Devices and methods of certain embodiments incorporatemechanically functional substrates comprising soft materials, forexample exhibiting flexibility and/or stretchability, such as polymericand/or elastomeric materials. In an embodiment, a mechanically matchedsubstrate has a modulus less than or equal to 100 MPa, and optionallyfor some embodiments less than or equal to 10 MPa, and optionally forsome embodiments, less than or equal to 1 MPa. In an embodiment, amechanically matched substrate has a thickness less than or equal to 0.5mm, and optionally for some embodiments, less than or equal to 1 cm, andoptionally for some embodiments, less than or equal to 3 mm. In anembodiment, a mechanically matched substrate has a bending stiffnessless than or equal to 1 nN m, optionally less than or equal to 0.5 nN m.

In some embodiments, a mechanically matched functional substrate ischaracterized by one or more mechanical properties and/or physicalproperties that are within a specified factor of the same parameter foran epidermal layer of the skin, such as a factor of 10 or a factor of 2.In an embodiment, for example, a functional substrate has a Young'sModulus or thickness that is within a factor of 20, or optionally forsome applications within a factor of 10, or optionally for someapplications within a factor of 2, of a tissue, such as an epidermallayer of the skin, at the interface with a device of the presentinvention. In an embodiment, a mechanically matched functional substratemay have a mass or modulus that is equal to or lower than that of skin.

In some embodiments, a functional substrate that is thermally matched toskin has a thermal mass small enough that deployment of the device doesnot result in a thermal load on the tissue, such as skin, or smallenough so as not to impact measurement and/or characterization of aphysiological parameter, such as a characteristic of a biological fluid(e.g. composition, rate of release, etc.). In some embodiments, forexample, a functional substrate that is thermally matched to skin has athermal mass low enough such that deployment on skin results in anincrease in temperature of less than or equal to 2 degrees Celsius, andoptionally for some applications less than or equal to 1 degree Celsius,and optionally for some applications less than or equal to 0.5 degreeCelsius, and optionally for some applications less than or equal to 0.1degree Celsius. In some embodiments, for example, a functional substratethat is thermally matched to skin has a thermal mass low enough that isdoes not significantly disrupt water loss from the skin, such asavoiding a change in water loss by a factor of 1.2 or greater.Therefore, the device does not substantially induce sweating orsignificantly disrupt transdermal water loss from the skin.

In an embodiment, the functional substrate may be at least partiallyhydrophilic and/or at least partially hydrophobic.

In an embodiment, the functional substrate may have a modulus less thanor equal to 100 MPa, or less than or equal to 50 MPa, or less than orequal to 10 MPa, or less than or equal to 100 kPa, or less than or equalto 80 kPa, or less than or equal to 50 kPa. Further, in someembodiments, the device may have a thickness less than or equal to 5 mm,or less than or equal to 2 mm, or less than or equal to 100 μm, or lessthan or equal to 50 μm, and a net bending stiffness less than or equalto 1 nN m, or less than or equal to 0.5 nN m, or less than or equal to0.2 nN m. For example, the device may have a net bending stiffnessselected from a range of 0.1 to 1 nN m, or 0.2 to 0.8 nN m, or 0.3 to0.7 nN m, or 0.4 to 0.6 nN m.

A “component” is used broadly to refer to an individual part of adevice.

“Sensing” refers to detecting the presence, absence, amount, magnitudeor intensity of a physical and/or chemical property, for example atissue parameter or an environmental parameter. Useful device componentsfor sensing include, but are not limited to electrode elements, chemicalor biological sensor elements, pH sensors, temperature sensors, strainsensors, mechanical sensors, position sensors, optical sensors andcapacitive sensors.

“Actuating” refers to stimulating, controlling, or otherwise affecting astructure, material, environment or device component, such as a tissueor an environment. Useful device components for actuating include, butare not limited to, electrode elements, electromagnetic radiationemitting elements, light emitting diodes, lasers, magnetic elements,acoustic elements, piezoelectric elements, chemical elements, biologicalelements, and heating elements.

The terms “directly and indirectly” describe the actions or physicalpositions of one component relative to another component. For example, acomponent that “directly” acts upon or touches another component does sowithout intervention from an intermediary. Contrarily, a component that“indirectly” acts upon or touches another component does so through anintermediary (e.g., a third component).

“Encapsulate” refers to the orientation of one structure such that it isat least partially, and in some cases completely, surrounded by one ormore other structures, such as a substrate, adhesive layer orencapsulating layer. “Partially encapsulated” refers to the orientationof one structure such that it is partially surrounded by one or moreother structures, for example, wherein 30%, or optionally 50%, oroptionally 90% of the external surface of the structure is surrounded byone or more structures. “Completely encapsulated” refers to theorientation of one structure such that it is completely surrounded byone or more other structures.

“Dielectric” refers to a non-conducting or insulating material.

“Polymer” refers to a macromolecule composed of repeating structuralunits connected by covalent chemical bonds or the polymerization productof one or more monomers, often characterized by a high molecular weight.The term polymer includes homopolymers, or polymers consistingessentially of a single repeating monomer subunit. The term polymer alsoincludes copolymers, or polymers consisting essentially of two or moremonomer subunits, such as random, block, alternating, segmented,grafted, tapered and other copolymers. Useful polymers include organicpolymers or inorganic polymers that may be in amorphous, semi-amorphous,crystalline or partially crystalline states. Crosslinked polymers havinglinked monomer chains are particularly useful for some applications.Polymers useable in the methods, devices and components disclosedinclude, but are not limited to, plastics, elastomers, thermoplasticelastomers, elastoplastics, thermoplastics and acrylates. Exemplarypolymers include, but are not limited to, acetal polymers, biodegradablepolymers, cellulosic polymers, fluoropolymers, nylons, polyacrylonitrilepolymers, polyamide-imide polymers, polyimides, polyarylates,polybenzimidazole, polybutylene, polycarbonate, polyesters,polyetherimide, polyethylene, polyethylene copolymers and modifiedpolyethylenes, polyketones, poly(methyl methacrylate),polymethylpentene, polyphenylene oxides and polyphenylene sulfides,polyphthalamide, polypropylene, polyurethanes, styrenic resins,sulfone-based resins, vinyl-based resins, rubber (including naturalrubber, styrene-butadiene, polybutadiene, neoprene, ethylene-propylene,butyl, nitrile, silicones), acrylic, nylon, polycarbonate, polyester,polyethylene, polypropylene, polystyrene, polyvinyl chloride, polyolefinor any combinations of these.

“Elastomer” refers to a polymeric material which can be stretched ordeformed and returned to its original shape without substantialpermanent deformation. Elastomers commonly undergo substantially elasticdeformations. Useful elastomers include those comprising polymers,copolymers, composite materials or mixtures of polymers and copolymers.Elastomeric layer refers to a layer comprising at least one elastomer.Elastomeric layers may also include dopants and other non-elastomericmaterials. Useful elastomers include, but are not limited to,thermoplastic elastomers, styrenic materials, olefinic materials,polyolefin, polyurethane thermoplastic elastomers, polyamides, syntheticrubbers, PDMS, polybutadiene, polyisobutylene,poly(styrene-butadiene-styrene), polyurethanes, polychloroprene andsilicones. Exemplary elastomers include, but are not limited to siliconcontaining polymers such as polysiloxanes including poly(dimethylsiloxane) (i.e. PDMS and h-PDMS), poly(methyl siloxane), partiallyalkylated poly(methyl siloxane), poly(alkyl methyl siloxane) andpoly(phenyl methyl siloxane), silicon modified elastomers, thermoplasticelastomers, styrenic materials, olefinic materials, polyolefin,polyurethane thermoplastic elastomers, polyamides, synthetic rubbers,polyisobutylene, poly(styrene-butadiene-styrene), polyurethanes,polychloroprene and silicones. In an embodiment, a polymer is anelastomer.

“Conformable” refers to a device, material or substrate which has abending stiffness that is sufficiently low to allow the device, materialor substrate to adopt a desired contour profile, for example a contourprofile allowing for conformal contact with a surface having a patternof relief features. In certain embodiments, a desired contour profile isthat of skin.

“Conformal contact” refers to contact established between a device and areceiving surface. In one aspect, conformal contact involves amacroscopic adaptation of one or more surfaces (e.g., contact surfaces)of a device to the overall shape of a surface. In another aspect,conformal contact involves a microscopic adaptation of one or moresurfaces (e.g., contact surfaces) of a device to a surface resulting inan intimate contact substantially free of voids. In an embodiment,conformal contact involves adaptation of a contact surface(s) of thedevice to a receiving surface(s) such that intimate contact is achieved,for example, wherein less than 20% of the surface area of a contactsurface of the device does not physically contact the receiving surface,or optionally less than 10% of a contact surface of the device does notphysically contact the receiving surface, or optionally less than 5% ofa contact surface of the device does not physically contact thereceiving surface. Photonic devices of certain aspects are capable ofestablishing conformal contact with internal and external tissue.Photonic devices of certain aspects are capable of establishingconformal contact with tissue surfaces characterized by a range ofsurface morphologies including planar, curved, contoured, macro-featuredand micro-featured surfaces and any combination of these. Photonicdevices of certain aspects are capable of establishing conformal contactwith tissue surfaces corresponding to tissue undergoing movement.

“Young's modulus” is a mechanical property of a material, device orlayer which refers to the ratio of stress to strain for a givensubstance. Young's modulus may be provided by the expression:

$\begin{matrix}{{E = {\frac{({stress})}{({strain})} = {\left( \frac{L_{0}}{\Delta \; L} \right)\left( \frac{F}{A} \right)}}},} & (I)\end{matrix}$

where E is Young's modulus, L₀ is the equilibrium length, ΔL is thelength change under the applied stress, F is the force applied, and A isthe area over which the force is applied. Young's modulus may also beexpressed in terms of Lame constants via the equation:

$\begin{matrix}{{E = \frac{\mu \left( {{3\; \lambda} + {2\; \mu}} \right)}{\lambda + \mu}},} & ({II})\end{matrix}$

where λ and μ are Lame constants. High Young's modulus (or “highmodulus”) and low Young's modulus (or “low modulus”) are relativedescriptors of the magnitude of Young's modulus in a given material,layer or device. In some embodiments, a high Young's modulus is largerthan a low Young's modulus, preferably about 10 times larger for someapplications, more preferably about 100 times larger for otherapplications, and even more preferably about 1000 times larger for yetother applications. In an embodiment, a low modulus layer has a Young'smodulus less than 100 MPa, optionally less than 10 MPa, and optionally aYoung's modulus selected from the range of 0.1 MPa to 50 MPa. In anembodiment, a high modulus layer has a Young's modulus greater than 100MPa, optionally greater than 10 GPa, and optionally a Young's modulusselected from the range of 1 GPa to 100 GPa. In an embodiment, a deviceof the invention has one or more components having a low Young'smodulus. In an embodiment, a device of the invention has an overall lowYoung's modulus.

“Low modulus” refers to materials having a Young's modulus less than orequal to 10 MPa, less than or equal to 5 MPa or less than or equal to 1MPa.

“Bending stiffness” is a mechanical property of a material, device orlayer describing the resistance of the material, device or layer to anapplied bending moment. Generally, bending stiffness is defined as theproduct of the modulus and area moment of inertia of the material,device or layer. A material having an inhomogeneous bending stiffnessmay optionally be described in terms of a “bulk” or “average” bendingstiffness for the entire layer of material.

“Tissue parameter” refers to a property of a tissue including a physicalproperty, physiological property, electronic property, optical propertyand/or chemical composition. Tissue parameter may refer to a surfaceproperty, a sub-surface property or a property of a material derivedfrom the tissue, such as a biological fluid. Tissue parameter may referto a parameter corresponding to an in vivo tissue such as temperature;hydration state; chemical composition of the tissue; chemicalcomposition of a fluid from said tissue; pH of a fluid from said tissue;the presence of absence of a biomarker; intensity of electromagneticradiation exposed to the tissue; wavelength of electromagnetic radiationexposed to the tissue; and amount of an environmental contaminantexposed to the tissue. Photonic devices of some embodiments are capableof generating a photonic response that corresponds to one or more tissueparameters.

“Environmental parameter” refers to a property of an environment of aphotonic device, such as a photonic device in conformal contact with atissue. Environment parameter may refer to a physical property,electronic property, optical property and/or chemical composition or anenvironment, such as an intensity of electromagnetic radiation exposedto the device; wavelengths of electromagnetic radiation exposed to thedevice; a chemical composition of an environmental component exposed tothe device; chemical composition of an environmental component exposedto the device; amount of an environmental contaminant exposed to thedevice; and/or chemical composition of an environmental contaminantexposed to the device. Photonic devices of some embodiments are capableof generating a photonic response that corresponds to one or moreenvironmental parameters.

“Photonic response” refers to a response generated by one or morephotonic structures of a photonic device of the invention. Photonicresponses may correspond to one or more parameters including tissueparameters and/or environmental parameters. In some embodiments, aphotonic response is an optical signal, such as a spatial and/ortemporal resolvable optical signal. In some embodiments, a photonicresponse is a measurable change in one or more of: (i) wavelengths oflight scattered, transmitted or emitted by said photonic structures;(ii) intensity of light scattered, transmitted or emitted by saidphotonic structures; (iii) spatial distribution of light scattered,transmitted or emitted by said photonic structures; (iv) phases of lightscattered, transmitted or emitted by said photonic structures; and/or(v) diffraction pattern of light scattered, transmitted or emitted bysaid photonic structures. Photonic responses useful in certainembodiments include, for example, a spectroscopic response, acolorimeteric response or fluorometric response.

FIG. 1A shows an embodiment of a photonic device (100) for interfacingwith a tissue in a biological environment, including for example atissue mounted device as shown in the Figure. The photonic device (100)comprises a flexible, or stretchable substrate (110), and one or morephotonic structures (120) supported by the substrate (110) forgenerating a photonic response corresponding to one or more tissueparameters or environmental parameters. In the embodiment shown in FIG.1A, the photonic structures (120) are provided in an array, such as apixelated two dimensional array. In this embodiment, the photonicstructures (120) are comprised of micro-, or nan-encapsulated structures(130) that encapsulate colorimetric and/or fluorometric indicators(140), for example, that provide a change in one or more opticalproperty in response to a change in a physical property, a physiologicalproperty or composition of the tissue (or a material derived from thetissue such as a biofluid) or a change in a physical property orcomposition of the environment of the device. As shown in this Figure,the substrate (110) is in conformal contact with a tissue surface (180)of a tissue (170). Optionally fluidic structures (150) are provided inthe substrate (110) to provide for fluid communication and/or transportof fluid from the tissue surface (180) to at least portion of thephotonic structures (120), in particular for some embodiments theencapsulated colorimetric and/or fluorometric indicators (140).Furthermore, additional device components (160) can be supported bysubstrate (160), such as wireless communication components includingantenna and near field communication device elements, opticalcomponents, electrodes and electrode arrays, and semiconductorstructures or devices. FIG. 1A also shows a, optical detector (190),such as a two dimensional detector, in optical communication with device(100) and capable of measuring the photonic response from said photonicstructures (120). Optical detector (190) may be a camera or otherimaging device, such as a camera on a mobile detect, capable ofspatially and temporally resolving the photonic response from

FIG. 1B shows an embodiment of a method of sensing one or more tissueparameters of a tissue of the subject or environmental parameters. Instep 1, a tissue (170) is provided and contacted with the photonicdevice (100), such that conformal contact is established with a surfaceof the tissue. In some embodiment, the photonic device is provided in inconformal contact with tissue selected from the group consisting of:skin, heart tissue, brain tissue, muscle tissue, nervous system tissue,vascular tissue, epithelial tissue, retina tissue, ear drum, tumortissue, or digestive system structures. In this embodiment, establishingconformal contact provides the device (and optionally the photonicstructures thereof) in physical contact, thermal communication, opticalcommunication, electrical communication, fluid communication or anycombination of these. In step 2, a photonic response corresponding toone or more tissue parameters or environmental parameters is generated,such as a photonic response comprising an a spatially and temporallyresolvable optical signal. In some embodiments, the photonic device(100) comprises a flexible or stretchable substrate (110); and one ormore photonic structures (120) supported by the substrate (110). In step3, a photonic response from the photonic device is detected using anoptical detector. In an embodiment, the method comprises detectingelectromagnetic radiation scattered or emitted by the one or morephotonic structures, thereby generating a detector signal. Optionally,the detecting step of the electromagnetic radiation is carried out usinga mobile electronic device. In an embodiment, the method furthercomprises analyzing the detector signal, thereby determining said one ormore tissue parameters or environmental parameters.

The invention can be further understood by the following non-limitingexamples.

Example 1: Epidermal Photonic Devices for Quantitative Imaging ofTemperature and Thermal Transport Characteristics of the Skin

Precision characterization of temperature and thermal transportproperties of the skin can yield important information of relevance toboth clinical medicine and basic research in skin physiology. Here, wedescribe an ultrathin, compliant skin-like, or ‘epidermal’, photonicdevice that combines colorimetric temperature indicators with wirelessstretchable electronics for precision thermal measurements when softlylaminated on the surface of the skin. The sensors exploit thermochromicliquid crystals (TLC) patterned into large-scale, pixelated arrays onthin elastomeric substrates, the electronics provide means forcontrolled, local heating by radio frequency (RF) signals. Algorithmsfor extracting patterns of color recorded from these devices with adigital camera, and computational tools for relating the results tounderlying thermal processes near the surface of the skin lendquantitative value to the resulting data. Application examples includenon-invasive spatial mapping of skin temperature with milli-Kelvinprecision and sub-millimeter spatial resolution. Demonstrations inreactive hyperemia assessments of blood flow and hydration analysisestablish relevance to cardiovascular health and skin care,respectively.

Spatio-temporal imaging of skin temperature offers experimental andinvestigational value for detection of breast cancers and othersyndromes, as an adjunctive screening tool to mammography.¹⁻³ Therequired milli-Kelvin levels of precision and milli-meter scaleresolution are most commonly achieved by use of sophisticated infrareddigital imaging cameras. Widespread adoption of such technology islimited, however, by high capital costs, motion artifacts, and inabilityfor use outside of clinical or laboratory settings. Other low costthermography techniques has been exploited much earlier, for potentialscreening of deep venous thrombosis⁴⁻⁷, breast cancer⁸⁻¹⁰, spinal rootsyndromes^(11,12), chronic back pain¹³ and even pulmonologicaldiagnostics.¹⁴ Recent work^(15,16) demonstrates that electronictemperature mapping devices can be constructed in ultrathin, soft andcompliant formats, sometimes referred to as ‘epidermal’ due to thesimilarity of their physical characteristics to those of the skinitself. These systems offer impressive capabilities that bypass manylimitations of infrared cameras, but provide only modest spatialresolution and imaging fidelity, limited by multiplexing systems neededto address large sensor arrays. Untethered, wireless operation alsodemands data transmission components and power sources. Otherstretchable smart skin devices that can monitor the vital health signalsof the wearer with unprecedented function and comfort have beeninvestigated intensively.¹⁷⁻²⁶ Here, we introduce a simple alternativethat combines colorimetric readout and RF actuation for precisionmapping of thermal characteristics of the skin. The sensors exploitthermochromic liquid crystals (TLC) patterned into large-scale,pixelated arrays on thin elastomeric substrates. Co-integration withelectronics provides a means for controlled, local heating by radiofrequency (RF) signals, to enable not only mapping of temperature butalso intrinsic thermal constitutive properties. Uniform layers of TLCsin water-impermeable, non-stretchable thick plastic sheaths, and withoutelectronics, have been explored for skin thermography,²⁷⁻²⁹ but withoutthe ability to conform sufficiently well to the curved, textured surfaceof the skin for accurate, reproducible measurements. Such devices alsofrustrate transepidermal water loss. They thermally load the skin, andcause irritation at the skin interface, thereby preventing reliable,accurate evaluation or use in continuous modes, over long periods oftime. Thermochromic textiles are available for cosmetic and fashionpurposes,³⁰⁻³² but their inability to maintain intimate contact with theskin and the limited capacity to use known thermochromic dyes forprecision temperature evaluation prevent their use in the sorts ofapplications envisioned here. The devices reported here not only avoidthese drawbacks, but they also allow precise measurement of thermalconductivity and thermal diffusivity through analysis of spatio-temporalimages obtained during operation of integrated RF components.Conventional digital cameras and RF transmission systems enablesimultaneous readout of thousands of pixels at resolutions that exceedthose needed to image temperature and thermal property variations on theskin. The epidermal format induces minimal perturbations on the naturalmechanical and thermal properties of the skin. Results presented in thefollowing establish the foundational aspects in materials, mechanics andthermal physics for both electronically active and passive epidermal TLC(e-TLC) devices, including algorithms for extracting precision,calibrated data from color digital images. Demonstrations in reactivehyperemia assessments of blood flow, as it relates to cardiovascularhealth, and hydration analysis, as it relates to skin-care, provide twoexamples of use in clinically meaningful tests.

The e-TLC thermal imagers use a multilayer design that includes (1) athin (20 μm) black elastomeric membrane as a mechanical support and anopaque background for accurate colorimetric evaluation of the TLCmaterials, (2) an array of dots of TLC (i.e. pixels, with 25 μmthicknesses, and diameters of either 250 or 500 μm, spaced by 250 or 500μm), with an optional interspersed array of dots with fixed colors (with25 μm thicknesses, diameters of 400 μm, spaced by 600 μm) forcalibration, both delivered to the surface of the black elastomer bytransfer printing, (3) a thin (30 μm) overcoat of a transparentelastomer for encapsulation and (4) optional electronics in thin,stretchable configurations mounted on the back surface for activefunctionality described subsequently (details appear in FIG. 8 andSupplementary Note 1). The TLC material consists of microencapsulatedchiral nematic liquid crystals. With increasing temperature, the phasevaries from crystalline solid to sematic, cholestoric and, finally,isotropic liquid, all over a range of a few degrees, dictated by thechemistry.^(33,34) In the cholestoric phase, light that reflects fromthe TLC pixels spans a narrow wavelength range defined by phase coherentinteractions with the liquid crystal assemblies. Increases intemperature decrease the pitch, thereby leading to blue-shifts in thepeak wavelengths of this reflected light. This behavior provides thebasis for colorimetric optical readout. Other phases have no chiralnematic orientation of molecular planes and thus do not yield any strongwavelength dependence to the reflection. The small sizes and largespacings of the TLC and calibration pixels, taken together with the lowmodulus, elastic properties of the substrate, encapsulation layer andelectronics, yield soft, compliant mechanics in the overall e-TLCsystem. These properties yield devices are well suited for mounting onthe skin.

FIG. 2a shows an e-TLC on the skin of the forearm when twisted andgently poked with a mildly heated rod. Low interfacial stresses thatfollow from the low effective modulus and small thickness of the deviceenable adequate adhesion through van der Waals interactions alone. Thecollapse of a free-standing device under its own weight, as in the rightframe, provides qualitative evidence of these mechanicalcharacteristics. FIG. 2b shows a pair of magnified images of e-TLCdevices, those on the bottom include interspersed color calibrationpixels consisting of red, green and blue dye in a non-toxic acrylic base(aqueous dispersion of organic pigment and acrylic polymer, Createx). Acompleted device of this latter type placed on the curved surface of theback of the hand appears in FIG. 2c . As previously mentioned, thebackside of the black elastomer substrate provides a mounting locationfor stretchable electronics. The image in FIG. 2d shows an example of ane-TLC device with a wireless system integrated in this way, for remotedelivery of controlled levels of heat. The folded configuration revealspart of the serpentine antenna structure (inset). An illustration ofthis system, in the form of three dimensional finite element analysis(3D-FEA), appears in FIG. 2e . The antenna captures incident radiofrequency (RF) energy to power a Joule heating element (inset, FIG. 2e). The result provides well-defined, localized increases in temperature,as revealed in the pattern of colors in the TLC pixels of FIG. 2f andthe infrared images of FIG. 2g . As described subsequently, the resultsfrom measurements under such conditions allow determination of thethermal conductivity and thermal diffusivity of the skin.

A key design goal is to produce e-TLC systems that induce minimalperturbations to the skin, thereby avoiding irritation, enhancingwearability and ensuring accurate measurement capabilities. Themechanical and thermal properties are particularly important in thiscontext. Experimental and theoretical studies of the former reveal lowmodulus, elastic characteristics over large ranges of strain. FIG. 3ashows the stress/strain responses of an e-TLC device under staticuniaxial testing. The results agree well with the predictions of 3D-FEA.In particular, the TLC pixels (˜221 MPa) and elastomeric substrate (˜131kPa) yield an effective modulus (˜152 kPa and 178 kPa from 3D-FEA andexperiment, respectively) that is only slightly larger (by 16-35%) thanthe intrinsic value associated with the bare elastomer, and iscomparable to that of the epidermis itself. The TLC pixels experienceultra-low strain (e.g., <2%) even under extreme stretching (e.g., 200%),as shown in FIG. 3b . Negligible deformations of the TLC pixels, asobserved in experiment and FEA (FIG. 3b ), allow approximations forsimple, but quantitatively accurate, analytical solutions of themechanics (see Supplementary Note 2 and FIG. 9a ). The thicknesses,bending stiffnesses, effective moduli and stretchability of thesedevices are 50 μm, 3.0 nN-m, 178 kPa and beyond 200%, respectively;these characteristics are superior than those of typical, commerciallyavailable TLC sheets (Hallcrest) whose corresponding properties of ˜125μm, 570,000 nN-m, 3.3 GPa and ˜5% (Hallcrest). The differences aresignificant, at a qualitative level of importance for deployment on theskin. In particular, the collective mechanical characteristics allowlargely unconstrained natural motions of the skin, including wrinklingand stretching even in challenging regions such as the knees and elbows.Addition of calibration pixels reduces the stretchability and increasesthe modulus (FIG. 9b ), but retain elastic strain levels (50%) thatexceed those that can be tolerated by the epidermis (linear response totensile strain up to 15%, nonlinear to 30%, and rupture at >30%³⁵).Incorporating a wireless electronic heating system further reduces theaccessible strain, but with an elastic stretchability of nearly 20%,which is useful for many applications (see FIG. 10).^(36,37) Althoughthe characteristics of the antenna change with mechanical deformation,experiments indicate that uniaxial stretching (up to 50%) does notdisrupt the overall function or the efficiency of power harvesting (seeFIG. 11); bending decreases the efficiency only slightly.

The thermal characteristics of the systems define the thermal load onthe skin, as well as the overall time response. For an active e-TLCdevice, the thermal mass per unit area is ˜7.7 mJ·cm²·K⁻¹ (SupplementaryNote 3). This value corresponds to an equivalent of skin thickness of˜20 μm, i.e. only 25% of the thickness of the epidermis itself.³⁸ Watervapor permeability test on e-TLC and Feverscan™ strip devices(Supplementary Note 4 and FIG. 12) has revealed that e-TLC devicesprovide minor moisture barrier for operation on skin. Decreasing thethickness of the device increases the water permeation, as expected (seeFIG. 12b ). Additional increases can be achieved by microstructuring,i.e. introducing arrays of holes or pores. The small thermal mass andhigh water permeability minimize changes in skin temperature andhydration level induced by presence of the device. Temperatures measuredwith an infrared camera on the forearm adjacent to an e-TLC and directlyunderneath it (FIG. 13a-c ) show minimal differences. The effects of thedevice on skin hydration (FIG. 13d-e ) are also small. A mounted 80 μmthick e-TLC on well hydrated skin (˜35) leads to a small percentageincrease in hydration (7.5%) after 3 hours. For an otherwise identicalset of testing conditions, the Feverscan™ strip led to a ˜100% increasein hydration. For monitoring of transient processes, the time responseof the system is important. With geometries and materials investigatedhere, the response time for an e-TLC device is dominated by thethickness and thermal properties of the black elastomer substrate.Transient measurements reveal response times of less than ˜30 ms(Supplementary Note 5), consistent with estimates developed usinganalytical models (FIG. 14). The intrinsic switching times for most TLCmaterials are ˜3-10 ms.³⁹⁻⁴²

Reflection mode spectroscopic characterization (Zeiss Axio Observer D1)of the steady-state response of the TLC material to changes intemperature between 32° C.-39° C. show expected behaviors, as in FIG. 4a. With proper calibration, described next, the temperature extractedfrom the hue and saturation values determined using a typical digitalcamera (Canon 5D Mark II) with the e-TLC device held at a nominallyconstant temperature exhibits a standard deviation of ˜30 mK over ameasurement time of 760 s. This value is comparable to that observedfrom temperature readings simultaneously determined with an infraredcamera (˜50 mK) (FIG. 4b ). The measurement precision is, then, at least±50 mK under these experimental conditions Equivalent temperaturesextracted from analysis of color recorded at the calibration pixels(red, green, blue) show fluctuations with similar magnitudes, assummarized in FIG. 4c . These observations suggest that the process ofimage capture and color analysis enables levels of precision that arecomparable to those of infrared cameras, not limited by the physics ofthe TLC. Detailed calibration plots and information on temperatureextraction appear in FIG. 15.

Analysis of hue/saturation/value data obtained from the digital camerarepresents the simplest and most straightforward analysis approach.Sophisticated algorithms based on computer vision techniques areadvantageous, however, not only for color determination but for fullpixelated analysis of complete e-TLC devices. FIG. 4d illustrates anexample of a process that exploits computer vision code (OpenCV), inwhich an image of an e-TLC device that consists of a 7×7 pixel arrayundergoes a set of color extraction and data transformation steps(details in Supplementary Note 6). A Gaussian filter first reduces noisethrough smoothing to yield a gray scale rendering for use with anadaptive threshold that compensates for illumination non-uniformities.The output is a binary mask containing value “1” at bright areas and “0”elsewhere. A two-step erode/dilate process eliminates small specklesthat arise from defects. A full list of contours can be extracted fromthis “clean” image, in which each contour bounds a single pixel in thearray. An enclosing circle function uses the contours as inputs todefine the pixel positions, for extraction of color information from theoriginal, unprocessed image. A typical calibration that relates hue andsaturation values extracted in this manner to temperature evaluated withan infrared camera appears in FIG. 4e . The biggest advantage of usinghue/saturation/value (HSV) color space instead of red/green/blue (RGB)is that the color information is encoded only in two (hue andsaturation), rather than three (red, green and blue) channels. These twovalues are comparatively resilient to changes in lightning levels sincethat information is stored separately in the value channel. Any possiblehue/saturation combination can be represented by a point in polarcoordinates where radial coordinate corresponds to saturation andangular one to hue. The positions of the calibration set are marked withthe dots painted with the corresponding hue. These points define thetemperature calibration surface by means of two dimensional linear fit.The results allow any hue/saturation combination to be assigned to atemperature value, as indicated in the plot using a color gradient.

Scaled use of this process is summarized in FIG. 4f . Here, a full e-TLCdevice on a portion of the wrist where near-surface veins are locatedreveals corresponding variations in temperature of the epidermis. Thehue values across the e-TLC yield three dimensional temperature contourplots that reflect the blood vessels with high spatial resolution (FIG.4g ). A direct comparison with temperature distributions measured in thesame region with an infrared camera (FIG. 4h ) exhibits excellentagreement. Plots of the temperature extracted from these two sets ofresults at the locations indicated by the dashed red lines in FIG. 4g,happear in FIG. 4i . These results suggest suitability of e-TLC systemsfor mapping of vascular distributions in applications such as screeningfor deep venous thrombosis, without the need for costly infrared camerasystems.

In such practical situations, the lighting conditions can stronglyaffect the precision and accuracy of the temperature determination.⁴³⁻⁴⁶In particular, the hue and saturation depend on the type of light sourceused for illumination. The color calibration pixels provide a means tocompensate for such effects, since their known colors are influenced bythe lighting in the same way as the TLC. As a result, it should bepossible to develop algorithms that account for shifts in the apparentcolors of these calibration pixels and yield a set of numericalcompensations that can restore their actual, known colors. Applying thesame compensations to the TLC pixels will serve as the basis for atemperature evaluation process that is independent of illuminationconditions, within some reasonable range. Effects of three differentlightning conditions appear in FIG. 5. Red, green and blue colorcalibration pixels, interspersed across the entire device, are presentin this active e-TLC sample. FIG. 5a presents an image of the device,with circles that indicate the positions of the TLC pixels. A Jouleheating element is present in the center region. Fluorescent, lightemitting diode (LED) and halogen (FIGS. 5c, 5d and 5e ) light sourcesprovide a range of practical examples. The corresponding temperaturecalibration data appear in FIG. 5b . The circles correspond to thehue/saturation values of TLC pixels recorded at different temperaturesto define calibration fits for specific light sources. The starsdelineate the effect of illumination on the colors of the calibrationpixels. Red, green and blue calibration pixels are located at ˜5°, ˜100°and ˜240°, respectively. Since these colors are known, data from themallow extraction of compensation factors for any given lightingcondition. Applying the results to measurements of TLC pixelsdramatically reduce the sensitivity of the temperature detection processto lightning source. FIG. 5f presents computed temperatures evaluatedalong lines that pass through the central region while the Joule elementis activated. The results are comparable for all three lighting sources.To demonstrate the importance of proper calibration, FIG. 5g summarizesdata that exploit the fluorescent temperature fit for all lightingconditions explored here. Significant discrepancies occur, as might beexpected due to the different color temperatures of the halogen and LEDsources. The resulting discrepancies in temperature readings arereflected not only in the temperature maxima, but also the temperatureprofiles, shapes and noise levels, which again emphasize the importanceof proper calibration and potential for compensation approaches.

As suggested by the active e-TLC results in FIG. 5, the local Jouleheating element enables additional measurement capabilities. Inparticular, spatial and temporal variations in temperature at locationsnear this heater can be used, with thermal models, to extract thethermal conductivity and diffusivity of the skin. Increases intemperature of a few ° C. can be sufficient for accurate evaluation. Thethermal conductivity (k) can be determined by comparing measured steadystate distributions in temperature to axis-symmetric thermal conductionmodels (see Supplementary Note 7). Calculations based on this modelsuggest spatial decays in temperature (T_(sensor-layer)) that vary as1/r (except the central sensor), which can be written as

$\begin{matrix}{{T_{{sensor} - {layer}} \approx {T_{\infty} + \frac{Q}{2\; \pi \; {kr}}}},} & (1)\end{matrix}$

where r is the distance from the heat source, Q is the heat generated bythe Joule heating element, and T_(∞) is the temperature of surroundingair. An example appears in FIG. 6a , with details in FIG. 16 a,b,e.Calibration can be performed through measurements of materials withknown properties (FIG. 6b ). FIG. 6c indicates excellent correspondencebetween thermal conductivity of the skin evaluated with an active e-TLCand hydration levels determined with a moisture meter (DelfinMoistureMeterSC) that relies on electrical impedance. The quantitativevalues of k fall within a range that is consistent literature valuesdetermined by subcutaneous thermocouples and high speed radiometeretc.⁴⁷ By simplifying the heating element as a point heat source turningon at time t=0, the transient temperature variation can be analyticallysolved as (see Supplementary Note 8)

$\begin{matrix}{{{T_{{Sensor} - {layer}}(t)} \approx {T_{\infty} + {\frac{Q}{2\; \pi \; {kr}}{{erfc}\left( \frac{r}{\sqrt{4\; \alpha \; t}} \right)}}}},} & (2)\end{matrix}$

where α is the thermal diffusivity of the skin, and erfc(x) is thecomplementary error function. Therefore, transient temperature dataassociated with activation or deactivation of the Joule heating elementcan be used to determine thermal diffusivity, α, as illustrated in FIG.6d (see FIG. 16 a,b,f). As with conductivity, the device can becalibrated using samples with known diffusivity (FIG. 6e ). Here, awireless active e-TLC system serves as the measurement vehicle. The timedependence of the temperature, rather than the absolute values, issufficient for extraction of diffusivity. The device operates atfrequencies of ˜2 GHz with maximum power inputs of ˜2.5 W/kg for thesubject of the studies described here (i.e. one third of the power limitrecommended by the Federal Communications Commission's guidelines). Thevalues also correspond closely to the hydration level, as shown in FIG.6f . As with k, the values of α are consistent with literature reportsbased on techniques such as opto-thermal measurement.⁴⁸ The values of kand a can be combined to yield the product of the density (φ and heatcapacity (c) of skin, based on the relation (cρ=k/α). The calculations(See FIG. 16g ) show that the heat capacity increases slightly with theincrease of hydration level (assuming that p is approximately constant),which is consistent with expectation since the heat capacity (˜4.2J/g/K) of water is larger than the human tissue (e.g., ˜3.7 J/g/K fordermis, ˜2.3 J/g/K for fat).⁴⁹

Spatio-temporal mapping even with passive e-TLC systems yields usefulinformation on blood circulation,^(50.51) maximal percentage increase inblood flow rate after occlusion,⁵² and duration of reactivehyperaemia.⁵³ Measurements of temperature fluctuations above the ulnarartery and adjacent veins serve as an important part of a reactivehyperaemia protocol. Here, the flow of blood is temporarily occluded bya pressure cuff on the upper arm, followed by abrupt release. FIGS.7A-7B summarize results of measurements performed with an e-TLC deviceand an infrared camera. FIG. 7C presents representative frames oftemperature distributions captured at 20 s intervals throughout theexperiment. Occlusion, which begins at t=0 s, causes the temperature ofthe skin above the ulnar artery and adjacent areas to decreasedrastically owing to lack of incoming blood flow and loss of heat to theenvironment. The minimum temperature is achieved at t=160 s; at thistime, the occlusion is released and blood flow resumes. Sharptemperature increases occur in areas above the blood vessels, as shownin FIG. 7C, until the temperature stabilizes. The responses of pixelsacross the array of the e-TLC vary widely depending on their distancefrom the blood vessels. The maximum temperature fluctuations are ˜1.2°C. and occur immediately above the ulnar artery; the minimum temperaturefluctuations are ˜0.4° C. and occur at locations away from near-surfaceblood vessels. Direct comparisons of spatio-temporal variations intemperature obtained from the e-TLC show quantitative agreement withresults from an infrared camera (FIG. 17). FIGS. 7D-7E highlighttemperature variations along horizontal and vertical lines illustratedin the right image of FIG. 7A. A thermal model of the human wrist(Supplementary Note 9 and FIG. 18) that accounts for both thetime-dynamic effect of occlusion and the thermal diffusion from theulnar artery can capture the effects revealed in the measurements (FIGS.7F-7G) and enable extraction of additional physiological information.The temporal variation of blood flow can be described with a piecewise,exponential type function,^(54,55) corresponding to the three stages ofthe process: pre-occlusion, vascular occlusion, and reperfusion. Theparameters characterizing this piecewise function can be determined byminimizing the average differences between the temperature-time profilespredicted by the model and those measured by the e-TLC device, duringeach stage. FIG. 7G shows that the calculated temperature history basedon the thermal model agrees with experiment at all six of the pixelsnear the artery (i.e., distance <6 mm). Due to simplifying assumptionsin the models, the FEA does not quantitatively capture the overshootbehavior observed in the two nearest sensors. Discrepancies at the twomost distant sensors can be attributed to the neglect of heatingassociated with a nearby vein (˜13 mm from the artery) in the model. Forvessel diameters and depths that lie within reported ranges(Supplementary Note 9), the peak blood flow velocity after occlusion iscalculated to be 58.8 cm/s, representing a three-fold increase over thebaseline of 19.6 cm/s, with reactive hyperemia duration of 144 s. Thesevalues match those reported in the literature for a person with lowcardiovascular risk.^(52,53)

The epidermal photonic systems, as embodied by the e-TLC devicesintroduced here, are useful for characterization of the skin and, byextension, important parameters relevant in determining cardiovascularhealth and physiological status. These same capabilities are also usefulin wound treatment and monitoring during a healing process, cancerscreening, core body temperature assessments and others of clinicalrelevance. In all cases, the ability to wear the devices continuously,over days or weeks, and to perform readout and power delivery via aconventional smartphone, represent uniquely enabling features for someembodiments. Photonic operation in the red and near infrared enable usein near-surface implantable diagnostics.

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Methods

Fabrication of e-TLC Thermal Imaging Devices.

The fabrication (details in FIG. 8) began with spin-coating and curing athin (20 μm) layer of poly(dimethylsiloxane) (PDMS, Sylgard 184, 40:1mixing ratio) mixed with Iron Oxide Pigment Black 11 (The Earth PigmentsCompany, LLC) on a substrate of poly(ethyleneterephlatate) (PET). A PDMSstamp with arrays of square posts (each post, 0.5 mm×0.5 mm over an areaof 15 cm²; see Supplementary Note 1a) was contacted against a layer ofmicroencapsulated thermochromic liquid crystals (Hallcrest SSN33R5W).Removing the stamp and drying it in air resulted in the formation of asolid layer of e-TLC material with an average thickness of 25 μm on theraised regions. A thermal release tape (Nitto Denko REVALPHA 90° C.)facilitated transfer of this material from the stamp to the surface ofthe black PDMS film. The device was completed by spin-coating and curinga thin (30 μm) layer of transparent PDMS on top of the structure, as anencapsulant. Fabrication of the wireless heater for the active e-TLCdevices began with spin-coating of a thin film of polyimide (SigmaAldrich) on a sacrificial layer of poly(methylmethacrylate) (PMMA; 100nm, MicroChem) on a silicon wafer. Metal-evaporation (Cr/Au, 5 nm/50nm), photolithography and wet-etching defined the serpentine structurefor the Joule heater. Additional polyimide spin-coating, oxygen reactiveion etching and metal deposition for contacts, interconnects, andantenna circuits completed the wireless system. Dissolving the PMMA andthen physically transferring the electronic structure to the back sideof the e-TLC device completed the fabrication.

Device Calibration and Test for Noise Level.

An e-TLC device was placed on a metal plate with black matt finish on ahotplate. Two white fluorescent light sources were placed on oppositesides of the device for illumination in a manner that avoided specularreflection. A digital camera (Canon Mark II 5D) and an infrared camera(FLIR ExaminIR) placed side-by-side were focused on the same area of thedevice at a distance of ˜30 cm. The angle between the cameras and eachof the light sources was ˜90 degrees. The device was heated to 40° C. onthe hotplate and then the hotplate was turned off. During the coolingprocess, high resolution images were collected every 10 seconds with thedigital camera; the infrared camera captured frames at a rate of 12.5s⁻¹. The process of cooling from 40° C. to 32° C. lasted about 20minutes. The color information of the TLC was extracted from 33° C. to39° C. with steps of 0.5° C. The set of algorithms developed toaccomplish this task are based on computer vision OpenCV(http://opencv.org/) library. The main functions are (in alphabeticorder) “adaptiveThreshold”, “cvtColor”, “dilate”, “drawContours”,“erode”, “findContours”, “GaussianBlur”, “getStructuringElement”,“imread”, “inRange”, “matchShapes”, “minEnclosingCircle”, “threshold”.In HSV color space, the light intensity information is stored in the“value” channel and is completely separated from the color informationwhich is encoded in the “hue” and the “saturation” channels. Hue andsaturation are, therefore, a natural basis for temperature calibrationsince they are not strongly affected by the change in illuminationintensity. Temperature calibration was constructed by means of twodimensional linear fit. The core function used in the process is “Istsq”from linear algebra module of Numerical Python (http://www.numpy.org/).Any combination of hue/saturation values can be assigned to atemperature value. Even for materials that are not temperature sensitivelike the calibration color pixels, their hue/saturation can be treatedas a specific temperature for consistency of analysis. To test the noiselevel and precision of the system, the hotplate temperature was set at afixed value; temporal fluctuations of TLC color, calibration dot colorand infrared emission were recorded using the two cameras over a periodof 15 minutes. The color changes were converted to temperaturefluctuation and compared to infrared fluctuation directly.

Reactive Hyperaemia Test.

A volunteer (female, 27 years old) reclined in a chair with her leftforearm secured gently to an arm rest using Velcro strips to reducemovement. A pressure cuff was secured around the subject's left bicep.An e-TLC device was placed on the skin of the left wrist approximatelyabove the ulnar artery. Applying puffs of compressed air ensured full,conformal contact. Infrared and digital cameras placed 30 cm above thesubject's left wrist were focused on the location of the device whileilluminated with white fluorescent lights. The subject was instructed torelax for 5 minutes. The cuff was inflated to a pressure of 250 mm Hgfor 160 seconds. Continuous high resolution color images and infraredtemperature measurements were then collected with the two cameras as theocclusion started and was then released. The total during of themeasurement period was 300 seconds.

Thermal Conductivity/Diffusivity and Hydration Measurements.

Thermal conductivity was determined by analyzing the spatialdistribution of temperature for a few seconds immediately afteractivation of a Joule heater in an active e-TLC device. To validate thecomputational models, an active e-TLC device was floated on the surfaceof a mixture of ethylene glycol/water preheated to ˜33° C. A constantvoltage supplied to the e-TLC Joule heating element created a steadystate temperature rise of a few degrees at the location of the heater.Images were then collected with a digital and infrared camera set upabove the device with only white fluorescent light sources. The spatialdecay of temperature in the e-TLC was recorded by analysis of imagesfrom the infrared camera and from color images of the device. The sameexperiment was performed on a volunteer's forearm skin. Here, differenthydration levels were achieved by applying various amounts of lotion tothe measurement location, prior to application of the active e-TLCdevice. Immediately after image capture, the e-TLC device was removedand a hydration meter was used to determine the actual moisture level(averaged from 5 readings). Measurements of thermal diffusivity used awireless, active e-TLC, with a transmission antenna located ˜10 cm awayand adjusted to achieve a peak change in temperature of a few degrees(RF power below 2.5 W/kg at frequencies between 1.95-2.35 GHz, tuned tomatch the response of the receiver antenna on the e-TLC). Both digitaland infrared cameras were focused on the device with a distance of 30cm. Videos with 60 second duration recorded the changes in temperatureassociated with activation and de-activation of the heater. Theexperiment was validated using the ethylene glycol/water system, andthen repeated on skin with different hydration levels, in proceduresotherwise similar to those for the thermal conductivity measurements.

Supplementary Note 1a: Fabrication Procedure for PDMS Post Stamp Usedfor Inking Liquid Crystal

1. Clean a 3″ Si wafer (Acetone, IPA->Dry 5 min at 110° C.).

2. Spin coat SU8 50 (microchem, 1000 rpm for 30 s, anneal 65° C. 10 min95° C. 30 min)

3. Pattern SU8 with 365 nm optical lithography through iron oxide mask(Karl Suss MJB3) develop in SU8 developer

4 post exposure bake at 65° C. 1 min 95° C. 10 min

5. STS ICP RIE silicon etch SF6 20 s at 20 w CF4 10 s at 0 w for 250cycles to achieve a hole depth of around 400 um

6. Mold the silicon template with PDMS

Supplementary Note 1b: Fabrication Procedure for a Single Heater withWired and Wireless Design

Prepare Polymer Base Layers

1. Clean a 3″ Si wafer (Acetone, IPA->Dry 5 min at 110° C.).

2. Spin coat with PMMA (poly(methyl methacrylate), spun at 3,000 rpm for30 s).

3. Anneal at 180° C. for 10 min.

4. Spin coat with polyimide (PI, poly(pyromelliticdianhydride-co-4,4′-oxydianiline), amic acid solution, Sigma-Aldrich,spun at 4,000 rpm for 30 s for wired design and 1,000 rpm for 30 s forwireless design).

5. Anneal at 110° C. for 30 s.

6. Anneal at 150° C. for 5 min.

7. Anneal at 250° C. under vacuum for 1 hr.

Deposit First Metallization

8. E-beam 5/50 nm Cr/Au.

9. Pattern photoresist (PR; Clariant AZ5214, 3000 rpm, 30 s) with 365 nmoptical lithography through iron oxide mask (Karl Suss MJB3).

Develop in aqueous base developer (MIF 327).

10. Etch Au with TFA Au etchant (Transene).

11. Etch Cr with CR-7 Cr Mask Etchant (Cyantek).

12. Remove PR w/Acetone, IPA rinse.

13. Dry 5 min at 150° C.

Isolate First Metallization and Pattern Via Holes

14. Spin coat with Pl.

15. Anneal at 110° C. for 30 s.

16. Anneal at 150° C. for 5 min.

17. Anneal at 250° C. under vacuum for 1 hr.

18. Pattern photoresist (PR; Clariant AZ4620, 3000 rpm, 30 s;) with 365nm optical lithography through iron oxide mask (Karl Suss MJB3). Developin aqueous base developer (AZ 400K, diluted 3:1).

19. Reactive ion etch (RIE; March CS-1701, 50 mTorr, 20 sccm O2, 150 W,35 min).

Deposit Second Metallization

20. E-beam 5/500 nm Cr/Au for wired design or 5/1600 nm Cr/Cu forwireless design.

21. Pattern PR AZ5214.

22. Etch Au with TFA Au etchant or etch Cu with TFA Cu etchant.cs

23. Etch Cr with Cr Mask Etchant.

24. Remove PR w/Acetone, IPA rinse.

25. Dry 5 min at 150° C.

Isolate Entire Device

26. Spin coat with Pl.

27. Anneal at 110° C. for 30 s.

28. Anneal at 150° C. for 5 min.

29. Anneal at 250° C. under vacuum for 1 hr.

30. Pattern PR AZ4620.

31. RIE (50 mTorr, 20 sccm O2, 150 W, 35 min for wired design and 120min for wireless design).

Release and Transfer

32. Release w/boiling Acetone.

33. Transfer to water soluble tape.

34. E-beam 3/30 nm Ti/SiO2.

35. Transfer to back of e-TLC device.

36. Bond thin, flexible cable (Elform, HST-9805-210) using hot iron withfirm pressure for wired heater

Supplementary Note 2: Analytic Solution of Spacing of e-TLC Dots DuringUniaxial Stretching

The deformation of an e-TLC device under uniaxial stretching (alonghorizontal direction) is analyzed to determine the change of spacingbetween pixels associated with the applied strain (ε). The e-TLCmaterial (˜221 MPa) is much stiffer than the elastomeric substrate (˜131kPa), and therefore undergoes negligible deformation, as evidenced bythe experiment images of FEA results in FIG. 3b . The stretchingdeformation is, as a result, mainly accommodated by the soft substratematerial. For pixels (in diameter of d_(TLC)) with an initial spacingΔ₀, the horizontal spacing (Δ_(horizontal)) after deformation is givenby

Δ_(horizontal)=Δ₀+(Δ₀ +d _(TLC))ε.  (S1)

The vertical spacing (Δ_(vertical)) decreases due to the Poisson effect.For sparsely distributed pixels (e.g., d_(TLC)<Δ₀), the mechanicalconstrains associated with the e-TLC on the transverse compression canbe neglected, such that the vertical spacing (Δ_(vertical)) afterdeformation can be approximated as

$\begin{matrix}{\Delta_{vertical} = {\frac{\Delta_{0} + d_{TLC}}{\sqrt{1 + ɛ}} - {d_{TLC}.}}} & \left( {S\; 2} \right)\end{matrix}$

Note that the transversely compressive strain of the soft substrate, dueto stretching (e), is given by ε_(compression)=1−(1+ε)^(−1/2), since itis nearly incompressible (i.e., Poisson ratio v=0.5). For Δ₀=0.3 mm,d_(TLC)=0.2 mm, as adopted in experiments, the analytic results in FIG.9a , based on Eqs. (S1) and (S2), agree well with the experiment and FEAresults.

Supplementary Note 3: Thermal Mass Calculation of e-TLC Device

The thermal mass of the devices are determined for 20 μm silicone andblack iron oxide substrate and 30 μm transparent silicone substrate. Thedevices have an overall aerial coverage of ˜15 cm². The calculatedthermal masses that follow are given as thermal mass per unit area ofskin. The device construction for the TCR device contains approximately8.7 ng·cm⁻² of Au, 56 μg·cm⁻² of PI, 55.8 μg·cm⁻² of Cu, 0.64 mg·cm⁻² ofblack iron oxide powder, 4.18 mg·cm⁻² of silicone substrate, ˜0.61mg·cm⁻² of liquid crystal materials (Hallcrest, density 0.97 g·cm⁻³).The material contributions to aerial thermal mass are: 21.48 μJ·cm⁻²·K⁻¹from Cu, 64.4 μJ·cm⁻²·K⁻¹ from PI, 0.42 mJ·cm⁻²·K⁻¹ from black ironoxide, ˜1.09 mJ·cm⁻²·K⁻¹ from liquid crystal (Hallcrest, specific heatSpecific heat 1.8 J·g⁻¹·K⁻¹), 6.11 mJ·cm⁻²·K⁻¹ from the silicone backing(calculate values) and negligible from Au. This results overall deviceaerial thermal masses of ˜7.7 mJ·cm⁻²·K⁻¹. The thermal mass of skindepends on the water content where thermal mass increases with skinhydration and water content². For hydrated skin, the heat capacity isapproximately 3.7 J·cm⁻³K-1, and the device aerial thermal mass of 7.7mJ·cm⁻²·K⁻¹ is equivalent to the aerial thermal mass of skin with athickness of 20.8 μm.

Supplementary Note 4: Water Vapor Permeability Test

Water permeability tests followed the ASTM E96-95 standard, and involvedevaluation of e-TLC devices (thicknesses of 80 μm, 50 μm and 30 μm) anda commercial Feverscan™ device (LCR Hallcrest; polyester covering film˜75 μm, liquid crystal layer ˜10-50 μm, black backing layer ˜10-20 μmand graphic print layer ˜10-20 μm). The experiments involved sealing thetops of identical jars, each containing a fixed amount of desiccant(˜97% anhydrous calcium sulfate and 3% cobalt chloride), with thedevices under test. Control samples consist of jars without any seal ontop. Diffusion of water vapor through the devices from the surroundingambient air causes increases in weight, due to uptake by the desiccant.All jars were placed in a room that has consistent temperature (˜22° C.)and humidity (˜50%). The weight gain of each jar was recorded at thesame time of day on a balance that has precision of 0.1 mg. By thistest, after a 4-day period, the weight of the jar sealed by theFeverscan™ remains unchanged, consistent with negligible waterpermeation. By contrast, weight of the jar with the 80 μm e-TLC deviceincreases by an amount that is nearly half (41%) of that compared to thecontrol. The 50 μm and 30 μm e-TLC devices exhibit weight increases thatare greater than half of the control, i.e. 60% and 62%, respectively.These results indicate that our formulation of PDMS, at the thicknessesused in our devices, provide only minor barriers to moisture,particularly when compared to conventional analogs.

Supplementary Note 5: Sensor Response Time

The TLC dot array is embedded in between two PDMS layers. The thicknessand thermal properties of the black PDMS substrate and the TLC layerwill both determine the heat transfer rate from the skin to the top ofTLC layer. The effect from the top encapsulation elastomer is neglectedto simplify the model.

A warm ethylene glycol bath heats up the entire device from the backsideof black PDMS substrate. The in-plane dimensions of the elastomer layerare much larger than its thickness such that the heat flux is mainlyalong the thickness direction, which can be represented by aone-dimensional heat transfer model described elsewhere.¹

The sensor response time is defined by the time at which the sensortemperature increase T_(sensor) reaches 90% of T₀. For 30 μm black PDMSand 25 μm TLC layer as used in the experiment, the response time ispredicted to be ˜30 ms. These agree reasonably well with theexperimentally measured sensor response time (for T_(sensor)=0.9T₀) of33 ms.

Supplementary Note 6: Color and Temperature Extraction Process

The only parts of TLS sensor that are temperature sensitive are theliquid crystal dots. Finding them in the image and separating from blackelastomer background is necessary first stage in temperature extractionprocess. This is a typical computer vision problem (OpenCVhttp://opencv.org/). The essential steps of the process are illustratedin FIG. 4a . First frame show the original picture of 7×7 area of thesensor array. Second is the output of Gaussian filter which reduce noisethrough image smoothing. Gray scale (third frame) format is requiredinput for adaptive threshold (fourth frame). Adaptive threshold is therobust algorithm that is aware of the illumination non-uniformity atdifferent parts of the image. The output is the binary mask containingvalue “1” at bright areas and “0” elsewhere. Small speckles from thedefects are visible here as well. They are removed with the two steperode/dilate process. Erode (fifth frame) shrink the white areas inframe four by removing few pixels at the border. Due to the small sizeof the defects they vanish completely. Dilate step (sixth frame) expandthe white regions back restoring area of interest by adding the sameamount of pixels removed in the previous step. List of contours can beextracted from this “clean” image (seventh frame). Every contour isenclosing a single temperature sensitive dot. The shape of the dot isclosely reminiscent to circle. The obvious choice for dots positiondetection is the OpenCV's “enclosing circle” function which take acontour as an input. Last frame is the superposition of the originalimage and the set of corresponding positions (red dots) and enclosingcircles (cyan rings).

Typical output of the digital camera is red-green-blue (RGB) color map.Intensities of all colors are affected by illumination conditions duringthe experiment. Converting to hue-saturation-value (HSV) color spacemake analysis more resilient to the change in lightning due to the factthat intensity now is encoded in value channel and color is in hue andsaturation channels. In order to track the color change only hue andsaturation are of interest. FIG. 4b show the calibration we use toconvert the colors into temperature. The dots plotted are positioned atcorresponding hue/saturation values and painted with their hue value.Background is the temperature evaluated from them with two dimensionallinear fit.

Supplementary Note 7: Steady-State Thermal Conduction Model forPrediction of Thermal Conductivity

A Cartesian coordinate system is set such that the origin is located atthe center of the top surface of PDMS, as shown in FIGS. 15a and 15b ,where the schematic illustrations of the device geometry, from both the3D and cross-sectional views, are presented. FEA indicates that theultrathin e-TLC dots (˜20 μm) have negligible effects on the temperaturedistributions, and thus are not considered in the analytic model. Theskin layer (homogenized from real skin and the underlying tissues, withthe thickness >2 mm) are usually much thicker than the PDMS layer (witha thickness of ˜60 μm), such that it can be considered as infinitelythick. The steady-state heat conduction equation is∂²T/∂x²+∂²T/∂y²+∂²T/∂z²=0 for both the PDMS and skin, where T is thetemperature. The square shaped resistor (a_(Resistor)×b_(Resistor))serves as the heat source, with the heat generation Q that pumps intothe PDMS and skin. This can be modeled as a surface heat flux(q₀=Q/(a_(Resistor)b_(Resistor))) for the bilayer system, i.e.,q₀=q_(zPDMS)|_(z=−H) _(PDMS) −g_(ZSkin)|_(z=−H) _(PDMS) for the regionoccupied by heat source. The free, top surface of the PDMS has naturalconvection with the surrounding air (T_(∞)), i.e.,q_(zPDMS)|_(z−0)=h(T−T_(∞)), with h denoting the heat transfercoefficient. The continuity conditions include [T]=0 and [q_(z)]=0across the PDMS/skin interface, where [ ]=0 stands for the jump acrossthe interface. By adopting the approach of double Fourier transform, thetemperature at the sensor plane (z=−H_(Sensor)) is obtained as

$\begin{matrix}{T_{{Sensor} - {layer}} = {T_{\infty} + {\frac{4\; q_{0}}{\pi^{2}k_{PDMS}} \cdot {\quad{{\int_{0}^{\infty}{{\cos \left( {\omega \; x} \right)}d\; \omega \; {\int_{0}^{\infty}\frac{\sin \; \frac{a_{Resistor}\omega}{2}\sin \frac{\; {b_{Resistor}\zeta}}{2}\left( {e^{\eta \; H_{Sensor}} + {\frac{{k_{PDMS}\eta} - h}{{k_{PDMS}\eta} + h}e^{\eta \; H_{Sensor}}}} \right){\cos \left( {\zeta \; y} \right)}d\; \zeta}{\omega \; \zeta \; {\eta \left\lbrack {{\left( {1 + \frac{k_{Skin}}{k_{PDMS}}} \right)e^{\eta \; H_{PDMS}}} - {\frac{{k_{PDMS}\eta} - h}{{k_{PDMS}\eta} + h}\left( {1 - \frac{k_{Skin}}{k_{PDMS}}} \right)e^{{- \eta}\; H_{PDMS}}}} \right\rbrack}}}}},}}}}} & \left( {S\; 3} \right)\end{matrix}$

where the subscripts ‘PDMS’ and ‘skin’ denote the PDMS and skin,respectively, k is the thermal conductivity. Eq. (S3) corresponds to thetemperature solution of the forward thermal conduction problem, giventhe thermal conductivity of the skin layer. The parameters adopted inexperiments include a_(Resister)=b_(Resister)=0.5 mm, h=5 W·m⁻²K⁻¹,H_(sensor)=30 μm, H_(PDMS)=60 μm, k_(PDMS)=0.16 W·m⁻¹K⁻¹, and thethermal diffusivity a_(PDMS)=1.07 m²·s⁻¹. For a representative value ofk_(skin)=0.31 W·m⁻¹K⁻¹ and Q=3.8 mW, the distribution of temperature atthe sensor plane, as given by Eq. (S3), is shown in FIG. 15c , whichagrees reasonably well with FEA results (FIG. 15d ). The temperatureprofile along the x axis (in FIG. 15e ) is in quantitative agreementwith the FEA results. The relatively large discrepancy at the centerregion is mainly attributed to the assumption of homogenious heatgeneration q₀ through the entire heater, adopted for the aim of modelsimplification. FIG. 15e also shows the temperature gradient is obviousin the region within a distance of ˜4 mm from the heater center. For thesensors far from the heater (0.5 by 0.5 mm), the temperaturedistribution can be approximated by the simple solution of a point heatsource, i.e.,

$\begin{matrix}{{T_{{Sensor} - {layer}} \approx {T_{\infty} + \frac{Q}{2\; \pi \; k_{Skin}r}}},} & \left( {S\; 4} \right)\end{matrix}$

where the ultrathin PDMS layer is neglected, and r=√{square root over(x²+y²)} is the in-plane distance from the origin. FIG. 15e demonstratesthat this approximate solution has very good accuracy forr≥a_(Resister)/2. This simplified solution is adopted to predict thethermal conductivity of skin by fitting the temperature data from thee-TLC device, as shown in FIG. 6a for an example with T_(∞)=33.9° C. andQ=3.83 mW. FIG. 6b demonstrates the prediction of thermal conductivityfor the calibration experiment, in which the water/ethylene glycolsolutions with different mixing ratios are adopted to mimic real skin indifferent hydration levels. The thermal conductivities predicted by thecurrent model agree fairly well with those reported in the literature(MEGlobal, Ethylene Glycol Product Guide).

Supplementary Note 8: Transient Thermal Conduction Model for Predictionof Thermal Diffusivity

To simplify the analyses for the transient thermal conduction problem,we continue to assume that the heater is a point heat source. Considerthat the heater is turned on at time t=0, the induced transienttemperature solution is given by

$\begin{matrix}{{{T_{{Sensor} - {layer}}(t)} \approx {T_{\infty} + {\frac{Q}{2\; \pi \; k_{skin}r}{{erfc}\left( \frac{r}{\sqrt{4\; \alpha_{skin}t}} \right)}}}},} & \left( {S\; 5} \right)\end{matrix}$

where α_(skin) is the thermal diffusivity of the skin, and erfc(x) isthe complementary error function. For the representative value ofk_(skin)=0.31 W·m⁻¹K⁻¹, α_(skin)=1.14 m²·s⁻¹, and Q=3.8 mW, the timedynamic temperature given by Eq. (S5) agree remarkably well with FEAresults, as shown in FIG. 15f , for three different points (with adistance of 0.5, 1.0 and 2.0 mm from the origin).

Based on Eq. (S5), we can determine the thermal diffusivity based on thetransient temperature data from the e-TLC device, even when the power isunknown (e.g., when the wireless system is adopted to power the heater).FIG. 6d gives an example of temperature profile at the sensor with adistance of 0.5 mm from the heater, where the analytic curve with thethermal diffusivity of 0.43×10⁻⁷ m²/s gives the best match with theexperimental data. FIG. 6e demonstrates the predictions of thermaldiffusivity for the calibration experiment, which agree reasonably wellwith those reported in the literature (MEGlobal, Ethylene Glycol ProductGuide).

Supplementary Note 9: Mathematical Modeling of Reactive Hyperemia

A two-dimensional (2D), transient, heat transfer model of human wristwas developed, which considers the various tissues surrounding the ulnarartery, and quantitatively characterizes the heat exchange between theblood flow and the surrounding tissues. FIGS. 17a and 17b show theschematic illustration of the tissue geometry, in which a circular crosssection is adopted for the wrist to simplify the analyses. The blood atbody temperature flows through the circular artery embedded in the fatlayer, heating the surrounding tissues. The heat exchange between theblood flow and the fat layer across the artery wall is described with aheat convection model², which assumes the exchanged heat flux (q) to beproportional to the blood flow rate, i.e.

$\begin{matrix}{{q = {\frac{\rho_{b}c_{pb}{\omega_{b}(t)}}{\pi \; D_{artery}}\left( {T_{body} - T_{s}} \right)}},} & \left( {S\; 6} \right)\end{matrix}$

where ρ_(b), c_(pb), ω_(b)(t) are the density, specific heat capacity,and time-dependent flow rate of the blood; D_(artery) is the diameter ofthe artery; T_(body) and T_(s) are the body temperature, and thetemperature of fat at the artery wall, respectively. Due to the heatingof the blood flow, the temperature distributes non-uniformly in thesetissues, which is governed by the temporal heat conduction equation of

${{\rho_{j}c_{j}\frac{\partial T_{j}}{\partial t}} = {{k_{j}\left( {\frac{\partial^{2}T_{j}}{\partial x^{2}} + \frac{\partial^{2}T_{j}}{\partial y^{2}} + \frac{\partial^{2}T_{j}}{\partial z^{2}}} \right)}\left( {j = {1\mspace{14mu} \ldots \mspace{14mu} 4}} \right)}},$

with the subscript representing different tissues (with skin as j=1, fatas j=2, muscle as j=3, and bone as j=4). The free, outer surface of theskin has natural convection with air, which usually cools down the skindue to a lower room temperature than body temperature. The interior bonelayer is assumed to maintain the core-temperature (close to the bodytemperature T_(boy)).

The modeling of occlusion involves two steps, starting from thesimulation of the steady-state heat conduction in the various tissuesdue to constant heating of blood flow, corresponding to the stage ofpre-occlusion (Stage 1). With the steady-state solution as an input, wefurther simulate the temporal changes in temperature distributions dueto the application and release of occlusion, corresponding to the stageof vascular occlusion (Stage II) and reperfusion (Stage III),respectively. Based on previous experimental data, the temporalvariation of blood flow during these different stages can be welldescribed by the following piecewise function^(2,3)

$\begin{matrix}{\mspace{79mu} {{{{\omega_{b}^{I}(t)} = \omega_{0}},{t \leq t_{{occ},{st}}}}\mspace{79mu} {{{\omega_{b}^{II}(t)} = {{\left( {\omega_{0} - \omega_{s}} \right){\exp \left( {{- t}\text{/}\tau_{0}} \right)}} + \omega_{s}}},{t_{{occ},{st}} < t \leq t_{{occ},{end}}}}{{\omega_{b}^{III}(t)} = \left\{ {\begin{matrix}{{{\left( {\omega_{\max} - \omega_{s}} \right){\sin^{2}\left\lbrack {{\pi \left( {t - t_{{occ},{end}}} \right)}\text{/}\left( {2\; t_{dw}} \right)} \right\rbrack}} + \omega_{s}},{t_{{occ},{end}} < t \leq \left( {t_{{occ},{end}} + t_{dw}} \right)}} \\{{{\left( {\omega_{\max} - \omega_{f}} \right){\exp \left\lbrack {{- \left( {t - t_{{occ},{end}} - t_{dw}} \right)}\text{/}\tau_{h}} \right\rbrack}} + \omega_{0}},{t > \left( {t_{{occ},{end}} + t_{dw}} \right)}}\end{matrix},} \right.}}} & \left( {S\; 7} \right)\end{matrix}$

where ω₀ represents the baseline blood flow; ω_(s) is the bloodperfusion after the occlusion is applied for a sufficiently long time,160 s in the case of experiments here; ω_(max) is the maximum hyperemicblood flow; τ₀ is a time constant depicting the falling speed of bloodflow after occlusion is applied; t_(dw) is the time required to reachthe maximum hyperemic blood flow after the release of occlusion; τ_(h)indicates the rate at which the blood flow returns to the baseline valueduring the reperfusion; t_(occ,st) and t_(occ,end) denote the startingand ending times of the occlusion, respectively. Except for t_(occ,st)and t_(occ,end), which are known in experiments (t_(occ,st)=0 s,t_(occ,end)=160 s), there are six parameters in this model of reactivehyperemia which can be varied to simulate the temperature history ofblood perfusion. The aim of the thermal analyses is to obtain anoptimized set of parameters that can minimize the average differencebetween the simulations and experiment data of temperature-time profileat those sensors with a distance≤7 mm from the artery (FIG. 7g ). Thebaseline blood flow ω₀ does not involve the occlusion process, andtherefore can be determined using the temperature value measured beforethe occlusion (Stage I). The blood flow ω_(s) and time parameter τ₀(only related to Stage II) are determined by the measuredtemperature-time profile during Stage II, and the other three parameters(ω_(max), t_(dw) and τ_(h)) are determined by the data during Stage III.In total, there are six parameters in our simulations, i.e., ω₀,α=ω_(s)/ω₀, β=ω_(max)/ω₀, τ₀, t_(dw) and τ_(h), whose ranges are listedin Supplementary Table 1, based on reported experiments^(2,3)

Finite element analyses (FEA) were adopted to solve the above transientheat transfer equation, and determine the temperature distributionnumerically. 4-node linear heat transfer elements were used, and refinedmeshes were adopted to ensure the accuracy. The boundary conditionsinclude the prescribed temperature (T=T_(body)) in the bone layer, theheat convection at the artery wall with blood flow of body temperature(i.e., Eq. (S6)), and the natural convection at the outer surface ofskin with air of room temperature (˜27.0° C.). The geometric andthermal-physical properties of various tissues are given inSupplementary Table 2. For the reactive hyperemia model described above,the baseline blood flow rate is determined as ω₀=30 mL/min (19.6 cm/sfor a vessel diameter of 1.8 mm), which could minimize the differencebetween FEA and experiment, i.e., the variance, as shown in FIG. 17c .Based on ω₀=30 mL/min, the calculated temperature decay from the arteryat the steady state indeed agree well with experiment data (FIG. 17d ).To minimize the temperature variance during stage II (FIG. S10 e), theblood flow ω_(s) and time parameter τ₀ are determined as ω_(s)=1.5mL/min and τ₀=2 s. Similarly, the other three parameters correspondingto stage III can be obtained as ω_(max)=90 mL/min (58.8 cm/s), t_(dw)=15s and τ_(h)=35 s. For this set of parameters, the temperature-timeprofile obtained from FEA agrees reasonably well with the experimentresults (FIG. 6g ) for all the sensor points close to the artery.

SUPPLEMENTARY TABLE 1 The parameter range in the model of reactivehyperemia for simulations. ω₀ (mL/ min) α = ω_(s)/ω₀ β = ω_(max)/ω₀τ₀(s) t_(dw) (s) τ_(h)(s) Range [10, 45] [0.05, 0.25] [3, 10] [2, 6][15, 45] [35, 75]

Supplementary Table 2. The geometric and thermal-physical properties ofvarious tissues for the wrist, where t denotes the thickness, D is thediameter of the artery, and d is the depth of the artery. Parameter SkinFat Muscle Bone Blood ρ (kg/m³)^((2,4)) 1085 850 1085 1357 1069 c_(ρ)(J/kg/K) ^((2,4)) 3680 2300 3768 1700 3659 k (W/m/K) ^((5,7)) 0.47 0.160.42 0.75 / t (mm) ⁽⁵⁻⁷⁾ 1.0 4.4 13.6 10.0 / D (mm) ⁽⁸⁾ / / / / 1.8 d(mm) ^((9,10)) / / / / 2.2

REFERENCES

-   1 Webb, R. C. et al. Ultrathin conformal devices for precise and    continuous thermal characterization of human skin. Nat. Mater. 12,    938, (2013).-   2 Deshpande, C. Thermal analysis of vascular reactivity MS thesis,    Texas A&M University, (2007).-   3 Akhtar, M. W., Kleis, S. J., Metcalfe, R. W. & Naghavi, M.    Sensitivity of digital thermal monitoring parameters to reactive    hyperemia. J. Biomech. Eng-T. Asme. 132, 051005, (2010)-   4 Fiala, D., Lomas, K. J. & Stohrer, M. A computer model of human    thermoregulation for a wide range of environmental conditions: The    passive system. J. App. Physiol. 87, 1957-1972 (1999).-   5 Song, W. J., Weinbaum, S., Jiji, L. M. & Lemons, D. A combined    macro and microvascular model for whole limb heat transfer. J.    Biomech. Eng-T. Asme. 110, 259-268 (1988).-   6 Sieg, P., Hakim, S. G., Bierwolf, S. & Hermes, D. Subcutaneous fat    layer in different donor regions used for harvesting microvascular    soft tissue flaps in slender and adipose patients. Int. J. Oral.    Max. Surg. 32, 544-547 (2003).-   7 Shen, H. et al. A genomewide scan for quantitative trait loci    underlying areal bone size variation in 451 Caucasian families. J.    Med. Genet. 43, 873-880 (2006).-   8 Shima, H., Ohno, K., Michi, K. I., Egawa, K. & Takiguchi, R. An    anatomical study on the forearm vascular system. J. Cranio. Maxill.    Surg. 24, 293-299(1996).-   9 McCartney, C. J. L., Xu, D., Constantinescu, C., Abbas, S. &    Chan, V. W. S. Ultrasound Examination of Peripheral Nerves in the    Forearm. Region. Anesth. Pain. M. 32, 434-439 (2007).-   10 Kathirgamanathan, A., French, J., Foxall, G. L., Hardman, J. G. &    Bedforth, N. M. Delineation of distal ulnar nerve anatomy using    ultrasound in volunteers to identify an optimum approach for neural    blockade. Eur. J. Anaesth. 26, 43-46 (2009).

Example 2: Stretchable, Wireless Sensors and Functional Substrates forEpidermal Characterization of Sweat

This Example introduces materials and architectures for ultrathin,stretchable wireless sensors that mount on functional elastomericsubstrates for epidermal analysis of biofluids. Measurement of thevolume and chemical properties of sweat via dielectric detection andcolorimetry demonstrates some capabilities. Here, inductively coupledsensors comprising LC resonators with capacitive electrodes showsystematic responses to sweat collected in microporous substrates.Interrogation occurs through external coils placed in physical proximityto the devices. The substrates allow spontaneous sweat collectionthrough capillary forces, without the need for complex microfluidichandling systems. Furthermore, colorimetric measurement modes arepossible in the same system by introducing indicator compounds into thedepths of the substrates, for sensing specific components (OH⁻, H⁺, Cu⁺,and Fe²+) in the sweat. The complete devices offer Young's moduli thatare similar to skin, thus allowing highly effective and reliable skinintegration without external fixtures. Experimental results demonstratevolumetric measurement of sweat with an accuracy of 0.06 μL/mm² withgood stability and low drift. Colorimetric responses to pH andconcentrations of various ions provide capabilities relevant to analysisof sweat. Similar materials and device designs can be used in monitoringother body fluids.

1. INTRODUCTION

Emerging wearable sensor technologies offer attractive solutions forcontinuous, personal health/wellness assessment,^([1,2]) forensicexamination^([3]) patient monitoring^([4,5]) and motionrecognition.^([1,7]) Recent advances in epidermal electronics^([8])provide classes of skin-mounted sensors and associated electronics inphysical formats that enable intimate, conformal contact with the skin.The soft, non-irritating nature of this contact yields an interface thatsimultaneously provides high precision, accurate measurement ofbiophysiological parameters, such as temperature,^([9])hydration,^([10]) strain,^([11]) and biopotential.^([12]) Such epidermalsensors are ultrathin, breathable and stretchable, with mechanical andthermal properties that closely match to the skin itself, to enableeffective skin integration with minimum constraints on naturalprocesses. The results provide unique capabilities in long-term,reliable health monitoring.

An important measurement mode in such devices may involve the analysisof body fluids (blood, interstitial fluid, sweat, saliva, and tear), togain insights into various aspects of physiological health.^([13-16])Such function in wearable sensors, generally, and epidermal electronicsin particular, is relatively unexplored. Existing devices either usecomplex microfluidic systems for sample handling^([17-20]) or involvepurely concentration-based measurement without sample collection andstorage, or access to parameters related to quantity and rate.^([21-23])In addition, mechanical fixtures, straps and/or tapes that are typicallyrequired to maintain contact of these devices with the skin do not lendthemselves well to continuous, long term monitoring withoutdiscomfort.^([24]) In the following, a set of materials and devicearchitectures that provide advanced capabilities in this area isreported. The key concept involves the use of functional soft substratesto serve as a means for microfluidic collection, analysis andpresentation to co-integrated electronic sensors and/or external camerasystems. The pores of these substrates spontaneously fill with bodyfluids that emerge from the skin, where they induce colorimetric changesin the substrate and alter the radio frequency characteristics ofintegrated electrical antenna structures. The results offer valuableinsights into the properties and volume of sweat, and theirrelationships to fluctuations in body temperature,^([25]) fluid andelectrolyte balance,^([26]) and disease state.^([27]) The devices alsoeliminate the need for direct skin-electrode contacts, therebyminimizing irritation that can be caused by contact between the skin andcertain metals,^([28]) while at the same time enabling repeated use of asingle device with minimal noise induced by motion artifacts. Thesensors exploit inductive coupling schemes, without on-chip detectioncircuits but with some potential for compatibility using near-fieldcommunication systems that are found increasingly in portable consumerelectronic devices. The entire sensing system offers flexible andstretchable mechanics, with form factors that approach those ofepidermal electronics.

2. RESULTS AND DISCUSSION

FIG. 19a shows images and schematic illustrations of a typical device(22×28 mm² for the surface area of the substrate, and 10×15 mm² for thedimension of the sensor) that includes an inductive coil and a planarcapacitor formed with interdigitated electrodes. The coil consists offour turns of thin copper traces in a filamentary serpentine design thataffords both flexibility and stretchability. The width of the trace is140 μm, and the lengths of the inner and outer turns are 4.8 and 9.5 mm,respectively. The electrodes consist of serpentine wires (50 μm inwidth) that have lengths between 6.5 to 8.4 mm, to form 9 digits with adigit-to-digit spacing of 600 μm. The dielectric properties of themicroporous supporting substrate strongly influence the capacitance ofthe structure.

In this way, the sweat sensor enables capacitive detection of the changeof the dielectric properties of the substrate as its pores fill withbiofluids (e.g. sweat). An external primary coil generates a timevarying electromagnetic field that induces a current flow within thesensor. The impedance of the sensor is then determined by the amount ofsweat within the substrate; this impedance influences that of theprimary coil placed in proximity to the device. The resonance frequency(f₀) of the sensor can be determined from the frequency of a phase dip(or a peak in the phase difference, A 6, obtained from the subtractionof the phase of the primary coil with and without the sensor underneath)in the phase-frequency spectrum of the primary coil.^([29-32]) Atmeasurement frequencies examined here (100 to 200 MHz), free watermolecules are under the influence of δ relaxation.^([33]) The responsesof the functional polymer substrates only involve contributions frominduced charges. The movement of the water molecules and dynamics of theinduced charges are sufficiently fast to respond to the externalelectromagnetic field. As a result, the combined dielectric propertiesof substrate and the sweat exhibit an invariant dielectric response overa wide range of frequencies (FIG. 23(a)). For present purposes, thefrequency-dependence in the dielectric properties of the substrate canbe ignored.

The sensor offers mechanical properties (elastic modulus≈80 kPa) similarto those of the skin.^([34]) The thickness of the substrate (1 mm),along with its lateral dimensions and porosity define the amount offluid that it can capture. The devices exhibit robust, elastic behaviorunder axial stretching (FIGS. 19b and 19c ) and other more complex modesof deformation (FIG. 19d ). Attachment of the sensor onto the skin (FIG.19e ) using a thin layer of commercial spray-on bandage as adhesiveleads to reversible skin/sensor bonding that can withstand significantextension and compression of the skin with sufficient mechanicalstrength to prevent delamination (FIGS. 19f and 19g ).

In vitro experiments involve slow introduction of 0.6 mL of buffersolution (phosphate buffered saline, Sigma-Aldrich Corporation, St.Louis, Mo, USA) onto the substrates with a syringe pump, over the courseof ≈40 minutes (FIG. 20d ). The resonance frequency of the sensor (f₀),as measured by the shift of the phase peak of a primary coil placed inproximity to the device (FIG. 20c ), decreases with increasing buffersolution content in the substrate. This response reflects increases inthe permittivity due to replacement of air with buffer solution in thepores of the substrate, leading to an increase in the capacitance of theinterdigitated electrodes associated with their proximity to thesubstrate. For a typical porous polyurethane (PUR) (PURpermittivity=7,^([35]) PUR substrate permittivity=1.42 at 0.93 porosityin air) (FIG. 20a ), f₀ shifts from 195.3 to 143.3 MHz in thisexperiment (FIG. 20d ). Drying of the sensor in air at room temperatureleads to the recovery of f₀, eventually to the original value (195.3MHz) over a period of ≈6 hours, indicating a reversible response withrelative insensitivity to residual salt that might remain in thesubstrate.

Assessment of performance with human subjects involves use of sensors oncellulose paper (CP) and silicone substrates attached to the arms of twovolunteers. Reference substrates made of the same materials with similarsizes placed in close proximity to the sensors provide means fordetermining the accuracy and establishing a calibrated response (FIG.20b ). The monitoring includes measuring the value of f₀ of the sensorsand the weight of the reference substrates every 5 min for a period of 2hours. The results indicate that f₀ is inversely proportional to theweight of the reference sensor, such that the response can be calibratedwith any two measured weights. The calibrated results closely followweight changes of 0.4 (FIG. 20e ) and 0.2 g (FIG. 20f ) in the referencesubstrates, corresponding to 0.4 and 0.2 mL of sweat over the sensingareas.

Dimensional changes associated with deformation of the skin or swellingof the device caused by sweat absorption could, conceivably, lead tochanges in f₀. Strain induced effects can be examined by biaxiallystretching a device and measuring f₀ at various states of deformation(FIG. 24(a)). The results show changes of only ≈0.9 MHz for biaxialstrains of 27% (FIG. 20g ) that are comparable to those caused by theabsorption of water (FIG. 24(b)). The modest changes in f₀ under biaxialstretching may be attributed to the symmetric design of the sensor coilas well as mutual compensation of the changes in lengths and spacings ofthe interdigitated electrodes. The effects of temperature are alsosmall. In particular, data indicate (FIG. 23(b)) that f₀ shifts from199.25 MHz to 196.63 MHz when the temperature is changed from 25 to 45°C. Finally, although the salinity and ionic content of the sweat maylead to changes in both conductivity and permittivity, experiments withbuffer solutions having various concentrations of sodium chloride (0 to4.5 g/L) reveal only small variations in f₀ (=0.6 MHz; FIG. 20h ).

The sensors exhibit excellent repeatability and are suitable forrepeated use. Multiple (i.e. five) measurements using sensors on CP andsilicone substrates serve as demonstrations. Between each measurement,immersion in water followed by drying on a hot plate regenerates thedevices. The changes in f₀ are repeatable for experiments that involveinjection of 0.6 mL buffer solution (FIGS. 20i and 23(c)). The averagechange in f₀ is 58.3 MHz with a standard deviation of 1.1 MHz for thesensor on CP; the corresponding values for silicone are 60.1 MHz and 3.6MHz, respectively. The changes in f₀ undergo different temporalpatterns, as might be expected due to the differences in chemistry,microstructure and pore geometry for these two substrates. Measurementsover 3 hours show no drifts in f₀ (FIG. 23(d)). The noise levels are<0.7 MHz; this parameter, together with an average change of f₀ of 58.3to 60.1 MHz for 0.6 mL buffer solution over a surface area of 22×28 mm²,suggests a measurement accuracy of ≈0.06 mL/mm².

The coil structures can be mounted onto various types of functionalsubstrates. Demonstrated examples include recycled cellulose sponge(RCS), polyurethane sponge (PUR), polyvinyl alcohol sponge (PVAS),cellulose paper (CP), and silicone sponge (FIG. 21a ). Cutting with ahot-wire device (PUR, silicone) or with a razor blade (other) yields theappropriate lateral dimensions and thicknesses. Spin-coated siliconefilms with accurately controlled thickness (≈10 μm; FIG. 21b ) enablestrong bonding between each of these functional substrates and thesensors through surface chemical functionalization, while preventingdirect contact between the sensors and the sweat. Relativecharacteristics in water absorption are also important to consider, asdescribed in the following.

The percentage gain in weight of the various porous materials afterimmersion in water defines their ability to hold fluids; the results are≈2300% (RCS), ≈1200% (PUR), ≈750% (PVAS), ≈350% (CP), and ≈1500%(silicone) (FIG. 21c ). These data, together with measured volumechanges yield the porosity levels: 0.97 (RCS), 0.93 (PUR), 0.90 (PVAS),0.83 (CP) and 0.95 (silicone) (FIG. 21d ). The water permeability can bedetermined from the capillary water absorption rate by combining Darcy'slaw^([36]) and the Hagen-Poiseuille equation.^([37]) Strips of thesubstrates (3 mm in width and 28 mm in length) are partially immersedinto water with red dye (3 mm under the water). A camera enablesmeasurements of changes in height of the absorbed water as a function oftime (FIG. 21e ). The CP material exhibits the fastest absorption rate(complete filling in ≈6 s), followed by the RCS (≈25 s). The PUR showsthe smallest rate, with an increase in height of 8.3 mm over 130 s.These rates depend strongly on the pore size and degree ofinterconnectedness and on the contact angle. The latter can bedetermined optically (FIG. 25(b)); the former can be obtained byscanning electron microscopy (FIG. 25(a)) or by calculation andmeasurement of the height of absorbed water at a long period of time(details in supporting information). The permeability of the fivesubstrates are 2.4 (RCS), 0.3 (PUR), 0.4 (PVAS), 8.7 (CP), and 8.6(silicone) μm² (FIG. 21f ).

In addition to dielectric response, absorption of water changes both thetransparency, due to index matching effects, and the overall dimensions,due to swelling (FIGS. 22a and 22c ). These effects can be used asadditional measurement parameters to complement the electrical datadescribed previously. The optical behavior can be illustrated by placinga sensor on a region of the skin with a temporary tattoo pattern.Continuous introduction of a buffer solution, up to a total of 0.6 mL,leads to increasing levels of transparency. Selected regions of theimages in FIG. 22a can be used to obtain RGB (red, green, and blue)intensities at different locations. The resulting data (FIG. 22b )indicate that the water content is inversely proportional to the ratioof the RGB intensity on the sensor and the skin. The water also induceschanges in the lateral dimensions. These changes can be measured byoptically tracking the displacements of an array of opaque dots (Cr, byelectron beam evaporation through a shadow mask) on the device (FIG. 22c). The results indicate a large displacement response to introduction of0.2 mL of the buffer solution (≈2.3 mm dot displacement), but withdiminishing response for an additional 0.4 mL (≈0.5 mm dotdisplacement). Nevertheless, these motions, which may be limited bymechanical constraints associated with mounting on the skin, might havesome utility in measuring sweat loss.

The substrates can be rendered more highly functional, from an opticalstandpoint, by introduction of chemicals or immobilized biomolecules.Resulting interactions with the sweat can be evaluated throughelectrical dielectric measurement or simply colorimetric detection. Forexample, silicone substrates doped with colorimetric indicators rendersensitivity to relevant biophysical/chemical parameters, such as pHvalues (FIG. 22d ), free copper concentrations (FIG. 26(a)), and ironconcentrations (FIG. 26(b)). To demonstrate pH detection, standardbuffer solutions with pH values from 4 to 9 are introduced into asubstrate that is dyed with a mixture of several different pH indicators(bromothymol blue, methyl red, methyl yellow, thymol blue, andphenolphthalein). These chemicals reversibly react with free —OH groupsand/or protons in the buffer solutions, leading to changes in absorptionspectra. Accordingly, the substrate undergoes a series of color changesthat reveal the pH values (FIG. 4d ). In addition, buffer solutions withcopper (FIG. 26(a)) and iron (FIG. 26(b)) at physiologicalconcentrations (0.8 to 1 mg/L) can also be detected using similarcolorimetric schemes. The intensities of individual colors (red, green,and blue) extracted from the images determine changes in analyteconcentrations (FIGS. 22e, 22f and 22g ). This type of strategy haspotential utility when used in combination with the sorts of wirelessschemes introduced here. For example, near field communication^([38])enabled devices such as cellphones also offer digital image capturecapabilities, for simultaneous colorimetric measurement.

3. CONCLUSIONS

The results presented here provide materials and design strategies forintegrating flexible and stretchable wireless sensors on functionalsubstrates. Demonstrated devices intimately mounted on the skin enablenon-invasive, wireless quantification of sweat loss as well ascolorimetric detection of sweat composition. Similar strategies can beused to develop sensors for monitoring a range of key parametersassociated not only with sweat but with other body fluids.

4. EXPERIMENTAL SECTION

To fabricate the device, a layer of polydimethylsiloxane (PDMS, 20 μmthick) is first spin-coated onto a glass slide (FIG. 27(a)). Curing thePDMS at 120° C. for 10 min and treating its surface with reactive ionetching (RIE) for 5 min (20 sccm O₂, 300 mTorr pressure, 150 W power)allows conformal spin-coating of a layer of polyimide (PI, 1 μm thick)on top. A bilayer of chrome (5 nm) and gold (200 nm) deposited byelectron beam (ebeam) evaporation is photolithographically patterned toform serpentine interdigitated electrodes (FIG. 27(b)). An additionalspin-coated PI (1 μm) layer electrically insulates the surfaces of theelectrode patterns, while selective regions on the PI layer are etchedby RIE for electrical contact between the electrode and serpentine coilsformed by patterning a layer of ebeam deposited copper (6 μm) (FIG.27(c)). The entire patterns are encapsulated by another spin-coated PIlayer (1 μm). Patterned RIE yields an open mesh layout, capable ofrelease onto the surface of a target substrate by use of water-solubletape (Aquasol ASWT-2, Aquasol Corporation, North Tonawanda, N.Y., USA).To prepare the functional substrates, a layer of uncured silicone (10 μmthick) is spin-coated onto a water soluble tape fixed on its edges to aglass slide by Scotch tape. Pre-curing the silicone at 150° C. for 1 mintransforms the liquid precursor into a tacky, soft solid (FIG. 27(e)).Placing the substrates on the silicone film with gentle pressure allowsthe partially cured film to crosslink with porous structures on thesurface. The silicone and the substrates are then fully cured at 120° C.to achieve robust bonding (FIG. 27(f)). The resulting structure isremoved from the glass, and rinsed with water to remove the watersoluble tape. Deposition of Ti/SiO₂ (5/60 nm) onto the exposed backsideof the sensor facilitates chemical bonding to the PDMS film on thefunctional substrates after UV ozone activation. Dissolving the watersoluble tape yields an integrated device with excellent levels ofmechanical stretchability and flexibility (FIG. 27(g) and FIG. 19b ).The functional substrates can be immersed into colorimetric indicators,followed by baking at 100° C. on a hotplate to dry the devices.

Five hydrophilic porous substrates serve as the sweat absorptionmaterials, including Whatman GB003 cellulose paper (GE Healthcare LifeSciences, Pittsburgh, Pa., USA), Scotch-Brite recycled cellulose sponge(3M Cooperation, St. Paul, Minn., USA), polyvinyl alcohol sponge(Perfect & Glory Enterprise Co., Ltd., Taipei), Kendall hydrophilicpolyurethane foam dressing (Covidien Inc., Mans-feld, MA, USA), andMepilex silicone foam dressing (Mölnlycke Health Care AB, Sweden). Forcolorimetric detection, a universal pH indicator (pH 2-10) (RiccaChemical, Arlington, Tex., USA) yields responses to buffer solutionswith well-defined pH (Sigma-Aldrich Corporation, St. Louis, Mo, USA).Colorimetric copper and iron ion detection is enabled by a copper colordisc test kit (CU-6, Hach Company, Loveland, Colorado, USA) and an ironcolor disc test kit (IR-8, Hach Company, Loveland, Colorado, USA), whilestandard stock solutions of copper and iron (Hach Company, Loveland,Colorado, USA) are diluted to achieve different ion concentrations.

The sensors can be integrated onto the skin. Briefly, spray bandage(Nexcare No Sting Liquid Bandage Spray, 3M Cooperation, St. Paul, Minn.,USA) is first applied onto the corresponding skin region. Evaporation ofthe solvent results in a tacky, water-permeable film that does notsignificantly influence the transdermal water loss from the skin andprovides sufficient adhesion to fix the sweat sensors onto the skin. Thesensor is then applied to the skin with continuous pressure over severalseconds. The bonding is reversible, but is sufficiently strong toaccommodate heavy sweating and shear forces.

The electrical responses of the sensors are evaluated using a HP 4291Aimpedance analyzer (Agilent Technologies, Santa Clara, Calif., USA) witha frequency range from 1 MHz to 1.8 GHz. The analyzer connects to aone-turn hand-wound copper primary coil whose resonance frequency issignificantly different from the sweat sensor. The coil is placed 2 mmaway from the sweat sensor during the measurement. However, smallvariations in the distance between the coil and the sweat sensor aretolerable, with negligible effects on the results. A xyz mechanicalstage and a rotational platform allow manual adjustment of the positionand orientation of the primary coil relative to the sweat sensor. Theprimary coil provides a time varying electromagnetic field that inducesalternating voltages in the sweat sensor. Changes of sweat contentwithin the substrate of the sensor lead to changes in the capacitance ofthe sweat sensor and its f₀. A syringe pump (KD Scientific Inc.,Holliston, Mass., USA) is used to deliver buffer solutions to thesensors during the in vitro experiments. The sweat sensors with a CPsubstrate and a silicone porous material are mounted on the arms of twovolunteers for 2 hour in vivo testing, with reference substrates of thesame materials and sizes placed in close proximity to the sweat sensors(FIG. 20b ). For the first hour, the volunteers exercise continuously togenerate sweat, and then stop to rest for the second hour. During themeasurement, the sweat sensors remain on the skin, while the referencesensors are peeled off every 5 min to record their weight using aprecise balance and reattached back to the same positions afterwards.

The absorbance values are estimated from the digital images by accessingthe RGB (red, green, blue) values of the selected regions on theexperimental images using ImageJ.^([39]) The average RGB values aredetermined from multiple pixels enclosed within a rectangular framedrawn by ImageJ with a plugin called, “measure RGB”. The Absorbance (A)defined as the negative log of the transmittance (I_(n)/I_(blank)), isthen calculated using the following formula:

A=−log(I _(n) /I _(blank))  (1)

in which I_(n) denotes the R, G or B values for the functionalsubstrates and I_(blank) the R, G, or B value for the background, bothobtained from the experimental images.

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Supporting Information

1. Methods for Determination of Weight Gain, Porosity, and Permeability

The percentage weight gain (W %) of the substrates can be obtained bymeasuring the weight of the materials in dry (W_(dry)) andwater-saturated (W_(sat)) states. Thus, W % can be expressed as.

$\begin{matrix}{{W\%} = {\frac{W_{sat} - W_{dry}}{W_{dry}} \times 100\%}} & (1)\end{matrix}$

The porosity (ϕ) of the materials is determined by the volume of pores(V_(pores)) to the total volume of the medium (V_(bulk)), is thusdefined by

$\begin{matrix}{\varphi = {\frac{V_{pores}}{V_{bulk}} = \frac{\left( {W_{sat} - W_{dry}} \right)/\rho_{water}}{{\left( {W_{sat} - W_{dry}} \right)/\rho_{water}} + {W_{dry}/\rho_{bulk}}}}} & (2)\end{matrix}$

where, ρ_(water) and ρ_(bulk) are the density of the water and thesubstrate materials, respectively.

To obtain the water permeability of the substrates, the Darcy law^([1]),which describes the water flow in porous materials, can be used. It isfound that the pressure gradient (∇P) that causes the water to flow inthe porous materials can be described by

$\begin{matrix}{{\nabla P} = {\frac{\mu}{K}q}} & (3)\end{matrix}$

where q is the volume average velocity (or flux), which representsdischarge per unit area, with units of length per time. The factor K isthe permeability of the material and p the viscosity of the water.Determination of ∇P typically involves an experimental setup containingtwo chambers with well-controlled pressures. An alternative method usesthe Hagen-Poiseuille equation^([2]) to determine ∇P by considering theporous materials as bundles of capillaries. As a result, the pressuregradient can be further expressed as:

$\begin{matrix}{{\nabla P} = {\frac{\Delta P}{L} = \frac{8\; \mu \; Q}{\pi \; R^{4}}}} & (4)\end{matrix}$

where ΔP is the pressure loss, L the length of the pipe, μ the dynamicviscosity, Q the volumetric flow rate (volume of fluid passing throughthe surface of the pipe per unit time), R the radius of the capillaries.Combing Eq. (3) and Eq. (4) yields

$\begin{matrix}{{\frac{\mu}{K}q} = \frac{8\; \mu \; Q}{\pi \; R^{4}}} & (5)\end{matrix}$

Here, Q/πR² represents the interstitial velocity of the flow, while qrepresents the superficial velocity of the flow. As a result, the ratiobetween Q/πR² and q is equivalent to the porosity of the materials

$\begin{matrix}{\varnothing = {\left( \frac{{Q/\pi}R^{2}}{q} \right).}} & \;\end{matrix}$

Thus, Eq. (5) can be further simplified as

$\begin{matrix}{R^{2} = \frac{8K}{\varphi}} & (6)\end{matrix}$

The linear momentum balance of the flow within a capillary tube can beexpressed as

$\begin{matrix}{\frac{2\; \sigma \; {\cos (\theta)}}{R} = {{\rho \; {gh}} + \frac{8\; \mu \; {hh}^{\prime}}{R^{2}} + {\rho \frac{d\left( {hh}^{\prime} \right)}{dt}}}} & (7)\end{matrix}$

where terms from left to right refer to the capillary pressure, thehydrostatic pressure, the viscous pressure loss, and the inertia terms,respectively. In Eq. (7), a is the surface tension of water, h is theheight of water in the capillary tube at time t, and 6 is the contactangle at the interface of the capillary tube and the water. As theporous materials may not have uniform R (especially for the porousmaterials with amorphous pores), such as RCS, PVAS, and CP in FIG.25(a), it is possible to replace R in Eq. (7) with a more general termR_(s), which represents the static radius of the porous materials andcan be obtained from the equilibrium height (h_(eq)) in the static case(height of the absorbed water in the porous materials when t reaches ∞).The static radius R_(s) can be calculated from

$\begin{matrix}{h_{eq} = \frac{2\; \sigma \; {\cos (\theta)}}{R_{s}\rho \; g}} & (8)\end{matrix}$

As a result, Eq. (7) can be further expressed as

$\begin{matrix}{\frac{2\; \sigma \; {\cos (\theta)}}{R_{s}} = {{\rho \; {gh}} + \frac{8\; \mu \; {hh}^{\prime}}{R_{s}^{2}} + {\rho \frac{d\left( {hh}^{\prime} \right)}{dt}}}} & (9)\end{matrix}$

by considering a flow regime where the influence of inertia as well asthe influence of gravity can be neglected^([3,4]). Thus, Eq. (8) can besimplified to

$\begin{matrix}{\frac{2\; \sigma \; {\cos (\theta)}}{R_{s}} = \frac{8\; \mu \; {hh}^{\prime}}{R_{s}^{2}}} & (10) \\{or} & \; \\{\frac{hdh}{dt} = \frac{\sigma \; {\cos (\theta)}}{4\; \mu}} & (11)\end{matrix}$

Solving this ordinary differential equation with the initial conditionh(0)=0 leads to the Lucas-Washburn equation [4].

$\begin{matrix}{h^{2} = {\frac{\sigma \; R_{s}{\cos (\theta)}}{2\; \mu}t}} & (12)\end{matrix}$

According to Eq. (6), Eq. (12) can be further expressed as

$\begin{matrix}{h^{2} = {{\frac{\sigma \; R_{s}{\cos (\theta)}}{2\; \mu}t} \cong {\frac{4\; \sigma \; {\cos (\theta)}}{\mu}\frac{K}{\varphi \; R_{s}}t}}} & (13)\end{matrix}$

As a result, the permittivity (K) can then be determined using thefollowing equation

$\begin{matrix}{K = \frac{h^{2}\mu \; \varphi \; R_{s}}{4\; \sigma \; {\cos (\theta)}t}} & (14)\end{matrix}$

where h, t, ϕ, and R_(s) of individual materials can all beexperimentally determined, as summarized in Table 3.

TABLE 3 Parameters of the porous materials used for functionalsubstrates ρ_(bulk) (g/ Materials W % cm³) ϕ h (m) t(s) R_(s) (m) K(μm²) RCS 2296 1.5 0.97 0.025 25 2.52E−5 2.43 PUR 1184 1.2 0.94 0.008130 1.52E−4 0.33 PVAS 746 1.2 0.90 0.013 130 6.34E−5 0.41 CP 332 1.50.83 0.025 6 2.50E−5 8.67 Silicone 1502 1.2 0.95 0.025 70 1.78E−4 8.57

2. Experiments for Determination of Weight Gain, Porosity, andPermeability

R_(s) can be determined from the h_(eq) measurement, in which 50 cmstrips of the porous materials are partially immersed into the water(approximately 1 cm strip in the water), while the heights of the waterin the strips after one day immersion are measured. As PUR and siliconehave more uniform pore sizes (FIG. 25(a)), their R_(s) can also bedetermined by measuring the radii of 10 pores in their SEM images andtaking the average numbers. The contact angle 8 can be measured throughthe analysis of images taken by a camera on the interface of water andthe porous materials (FIG. 25(b)). The relation between h and t can beobtained using video captured throughout the process of waterabsorption.

REFERENCES

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Example 3: Epidermal Microfluidic Sweat Patch

This Example discloses an epidermal microfluidic sweat patchincorporating at least one microfluidic channel and a plurality ofcolorimetric indicators disposed within cavities of the patch. The patchoptionally includes a near-field communication coil.

Table 4 shows concentrations of parameters and chemical species relevantto sweat monitoring.

TABLE 4 Parameters and chemical species relevant to sweat monitoring.Median Constituents Concentration Range Sweat gland density 100pores/cm² 50~300 pores/cm² Sweat rate 50 μL/hour · cm² 12-120 μL/hour ·cm² pH 4.0-6.8 Glucose 0.17 mM 5.6 μM-2.2 mM Lactic acid   14 mM 3.7-50mM Chloride   23 mM 0.02-280 mM Sodium ion   31 mM 0.11-390 mM

FIG. 29 shows an exploded view of a colorimetric sensor comprising anear-field communication coil. FIG. 29 is a photograph of the device ofFIG. 28 adhered to the skin of a subject.

FIG. 30 illustrates a fabrication method for a sweat patch and anadhesion test on skin.

FIG. 31 illustrates an artificial sweat pore test using a syringe tofeed artificial sweat at a rate of 12 μL/hr.

FIG. 32 shows a sweat patch incorporating colorimetric detection ofvarious biomarkers for self-monitoring and early diagnosis. For example,FIG. 33 shows an absorbance spectrum illustrating the color change of areactant that may be used to determine sweat volume and rate. FIG. 34shows an absorbance spectrum and legend illustrating the color change ofa reactant(s) that may be used to determine sweat pH, which may becorrelated with sodium concentration, indicating to a user the propertime to hydrate. FIG. 35 shows an absorbance spectrum and legendillustrating the color change of a reactant(s) that may be used todetermine glucose concentration in sweat, which may be correlated withblood glucose concentration. FIG. 36 shows an absorbance spectrum andlegend illustrating the color change of a reactant(s) that may be usedto determine lactate concentration in sweat, which may provide anindication of shock, hypoxia and/or exercise intolerance.

As shown in FIG. 37, a sweat sensor incorporating colorimetric biomarkerindicators provides qualitative and quantitative data that may beobserved by the naked eye and/or wirelessly observed by a detectiondevice, such as a smartphone.

Example 4: Sweat Patches

Overview

Provided herein are epidermal microfluidic sweat patches for daily wearas personal healthcare monitoring systems that are highly conformableand stretchable. The patches allow for the non-invasive determination ofsweat rate, sweat volume, and biomarker concentration, thereby providingclinically reliable information. This technology relates toself-diagnostic systems for monitoring an individual's health state bytracking color changes of indicators within the devices by the naked eyeor with a portable electronic device (e.g., a smartphone). By monitoringchanges over time or trends, the disclosed devices may provide earlyindications of abnormal conditions.

The disclosed sweat sensor enables detection of sweat volume and rate,as well as concentration of biomarkers in sweat (e.g., pH, glucose,lactate, chloride, creatinine and ethanol) via various quantitativecolorimetric assays. In an embodiment, the colorimetric indicators areincorporated into a polydimethysiloxane (PDMS) substrate because PDMS isa silicon-based organic polymer approved for a wide range of medicalapplications, including contact lenses and medical devices.

Epidermal Microfluidics

Microfluidic analytical devices for sweat monitoring were developedbased on a 2D channel system within poly(dimethylsiloxane) (PDMS)without pumps, valves, or fluid detectors. The chemical and physicalcharacteristics of PDMS made it suitable for epidermal applications. Forexample, PDMS is optically transparent, elastomeric, nontoxic,chemically inert toward most reagents, and possesses a low surfaceenergy. The fabricated epidermal sweat patch was composed of fourindividual quantitative colorimetric detection reservoirs and anorbicular outer-circle serpentine fluidic channel (FIG. 39A). Each ofthe biomarker detection reservoirs holds 4 μL while the orbicular waterdetection channel contains 24 μL. The sample inlet located at the bottomof the device (0.5 cm²) may cover about 50 sweat glands, thusintroducing sweat into the device, filling the detection reservoirs, andallowing sweat to flow through the outer-circle channel forapproximately 6 hours calculated based on an average sweat rate of 12μL/hour·cm² for humans. Due to the interfacial permeability of PDMS,which is impermeable to liquid water but permeable to gases, the waterloss of the sweat patch was moderate (3% of the total volume during thesensor life-time). The device was 3 cm in diameter and 500 μm inthickness constructed with PDMS consisting of 30:1 (v/v) base:curingagent resulting in a modulus of 145 kPa. The mass of the device was ˜970mg.

The epidermal microfluidic sweat sensors were fabricated using softlithography. The schematic illustration and fabrication processes areshown in FIG. 38. A master device was prepared from a silicon wafer byphotolithography and dip-etching to generate a reverse image having 300μm deep channels. To produce replicas, the mixture of 30:1 (v/v)base:curing agent of PDMS was poured over the master that was coatedwith a thin layer of poly(methyl methacrylate) (PMMA) and cured at 70°C. for 1 h. Once the PDMS was fully cured, the replica was released fromthe master. The prepared replica was then sealed with a PDMS film byoxygen plasma bonding for 1 min to activate surface silanol groups toform siloxane bonds. Finally, the fabricated microfluidic devices wereattached to a commercial medical dressing (i.e., Tegederm®) via oxygenplasma bonding and applied on the skin surface. This epidermalmicrofluidic sweat-monitoring device was able to withstand significanttension, compression, and twist of the skin while maintaining sufficientadhesion (FIG. 38C).

Quantitative Colorimetric Detection of Biomarkers

The colorimetric determination holds great advantages for diagnosis inquantitative analysis. In this sweat sensor, four colorimetric analyseswere introduced for biomarkers being able to self-diagnosis and monitora variety of medical conditions. Each detection reservoir represented adifferent analyte for determination of (1) water (for sweat volume andrate evaluation), (2) pH, (3) glucose, (4) lactate, and (5) chlorideconcentrations.

Thermal regulation and dehydration are highly related to sweat rate andvolume and thus continuous monitoring is a vital tool for assessinghealth states of individuals and providing information relating toelectrolyte balance and rehydration. The orbicular channel in the sweatsensor was coated with cobalt (II) chloride (i.e., CoCl₂) contained in apolyhydroxyethylmethacrylate hydrogel (pHEMA) matrix. As the sweat isintroduced into the channel the blue colored anhydrous cobalt (II)chloride reacts with water turning into hexahydrate cobalt chloride(i.e., CoCl₂.6H₂O) presenting a pale purple color (FIG. 40A). Bydetermining the distance of color change within the channel during acertain period of time, the sweat rate and volume could be assessed.

Not only physical sweat analysis, but chemical detection of biomarkersin sweat is essential. In some embodiments, quantitative colorimetricassays were demonstrated with paper-based reservoirs individuallylocated in the middle of the sweat sensor. Filter paper was chosen as amatrix material among other materials (e.g., hydrogel, sol-gel, andagarose gel) since the hydrophilic cellulose fibers wicked biofluids ata fast absorption rate, as well as provided a solid support for assayreagent and allowed clear contrast regarding color changes.² Acolorimetric sweat sensor was developed that consisted of four biomarkerdetection reservoirs: pH, glucose, lactate, and chloride.

The pH value of sweat has been known to exhibit a proportionalrelationship with sweat rate and sodium ion concentration. As anindicator of proper hydration time for a user, sweat pH was determinedusing a universal pH indicator consisting of various pH dyes (e.g.,bromothymol blue, methyl red, and phenolphthalein), which covers a widerange of pH values. While the sweat was introduced in the reservoir, thepH indicator changed color based on the ratio of weak acid and itsconjugate base form of the indicator based on the Henderson-Hasselbalchequation. The color change was observed according to various pH valuesof buffer solution in a medically reliable range (i.e., pH 4.0-7.0) asshown in FIG. 40B and its respective spectrum is presented in FIG. 40C.

Glucose concentration in the sweat is one of the most biomarkers formonitoring health state, especially playing a crucial role for improvingdiabetes treatment. In this device, the glucose was detected based on anenzymatic reaction that governed the selectivity of the measurement.Physically immobilized glucose oxidase produced hydrogen peroxideassociated with oxidation of glucose and reduction of oxygen, next,iodide was oxidized to iodine by peroxidase, which was also contained inthe paper-based reservoir.³ Therefore, a color change was observed fromyellow to brown, the respective colors of iodide and iodine, to indicatethe concentration of glucose.³ The color change illustrating the glucoseconcentration is presented in FIG. 40B as well as the respectivespectrum in FIG. 40D. Thus, this device may warn of abnormal bloodglucose concentrations for not only diabetes patients but alsoprediabetes and healthy persons by correlating perspiration glucoseconcentration in a completely noninvasive manner on a daily basis.⁴

The sweat lactate concentration is an indicator of exercise intolerance,tissue hypoxia, pressure ischemia, and even pathological conditions(e.g., cancer, diabetes, and lactate acidosis).⁵ Lactate is produced byanaerobic energy metabolism from the eccrine gland, so lactateconcentration in perspiration is a good criterion for determiningindividuals' abilities to endure rigorous exercise, especially forathletes and military personnel, and/or severe physical activity whileon life support.⁶ Enzymatic reactions between lactate and co-factor NAD⁺by lactate dehydrogenase and diaphorase allowed a color change of achromogenic reagent (i.e., Formazan dyes) resulting in an orange color.As shown in FIGS. 22B and 22E, the color change within the detectionreservoir was observed with regard to the concentration of lactatewithin the medically relevant range of 1.5-100 mM.

The representative sweat tests rely on determination of chloride ionconcentration in perspiration. These tests may diagnosis cystic fibrosis(CF) since excreted chloride content increases when there are defectivechloride channels in sweat glands.⁷ Additionally, the level of chlorideis considered to be an index of hydration. Accordingly, the level ofchloride in sweat was determined using colorimetric detection bycompetitive binding between Hg²+ and Fe²⁺ with2,4,6-tris(2-pyridiyl)-s-triazine (TPTZ). In the presence of chlorideion, iron ion prefers to bind with TPTZ while Hg²⁺ participates asHgCl₂, which results in a color change from transparent to blue bindingwith respective metal ions. The quantitative colorimetric results areshown in FIGS. 40B and 40F.

Not only the biomarkers mentioned above, but copper ion, iron ion, andethanol concentrations in sweat may also be detected by colorimetricassay. The trace copper ion in sweat was determined using a1,2-bicinchoninate acid (BCA). The copper complex with BCA exhibited anintense purple color demonstrating a quantitative color change from 0 to1 mg/mL.⁸ Similarly, iron ions were detected by a colored complex formedwith 1,10-phenanthroline in the range of 0-0.8 mg/L.^(8b) Additionally,colorimetric detection of ethanol was demonstrated using an enzymaticreaction consisting of alcohol dehydrogenase, peroxidase, and formazandye.

Collectively, these quantitative colorimetric analyses providepre-diagnostic information of multiple biomarkers in sweat. By combiningthe colorimetric devices with telemedicine technology, this sweat patchcould provide a user-friendly self-monitoring system for daily wear.

Telemedicine Technologies

In order to provide personalized clinical health care with a smartphone,near field communication (NFC) electronics were applied to the sweatpatch. The NFC communication devices were fabricated with an ultrathinconstruction using ultralow modulus materials, which enable wirelesscommunication under extreme deformations in daily usage.⁹ The NFC coilswere incorporated on the sweat patch as shown in FIG. 41A. Thebiomedical information of sweat is quantitatively analyzed by takingimages of the sweat sensor showing the color changes of the reservoirs(FIG. 41B). Using wireless NFC electronics to communicate to asmartphone permits the images to be examined based on an RGB digitalcolor specification, converted into health informatics (e.g.,concentration of biomarkers) and optionally transmitted from anindividual's smartphone to medical staff or a medical records database.

REFERENCES

-   1. McDonald, J. C.; Whitesides, G. M., Poly(dimethylsiloxane) as a    Material for Fabricating Microfluidic Devices. Accounts of Chemical    Research 2002, (7), 491-499.-   2. Martinez, A. W.; Phillips, S. T.; Whitesides, G. M.; Carrilho,    E., Diagnostics for the Developing World: Microfluidic Paper-Based    Analytical Devices. Analytical Chemistry 2010, 82 (1), 3-10.-   3. (a) Martinez, A. W.; Phillips, S. T.; Butte, M. J.;    Whitesides, G. M., Patterned Paper as a Platform for Inexpensive,    Low-Volume, Portable Bioassays. Angewandte Chemie International    Edition 2007, 46 (8), 1318-1320; (b) Martinez, A. W.; Phillips, S.    T.; Carrilho, E.; Thomas, S. W.; Sindi, H.; Whitesides, G. M.,    Simple Telemedicine for Developing Regions: Camera Phones and    Paper-Based Microfluidic Devices for Real-Time, Off-Site Diagnosis.    Analytical Chemistry 2008, 80 (10), 3699-3707.-   4. Moyer, J.; Wilson, D.; Finkelshtein, I.; Wong, B.; Potts, R.,    Correlation Between Sweat Glucose and Blood Glucose in Subjects with    Diabetes. Diabetes Technology & Therapeutics 2012, 14 (5), 398-402.-   5. (a) Polliack, A.; Taylor, R.; Bader, D., Sweat analysis following    pressure ischaemia in a group of debilitated subjects. J Rehabil Res    Dev 1997, 34 (3), 303-308; (b) Biagi, S.; Ghimenti, S.; Onor, M.;    Bramanti, E., Simultaneous determination of lactate and pyruvate in    human sweat using reversed-phase high-performance liquid    chromatography: a noninvasive approach. Biomedical Chromatography    2012, 26 (11), 1408-1415.-   6. Jia, W.; Bandodkar, A. J.; Valdes-Ramirez, G.; Windmiller, J. R.;    Yang, Z.; Ramirez, J.; Chan, G.; Wang, J., Electrochemical tattoo    biosensors for real-time noninvasive lactate monitoring in human    perspiration. Anal Chem 2013, 85 (14), 6553-60.-   7. Mishra, A.; Greaves, R.; Massie, J., The Relevance of Sweat    Testing for the Diagnosis of Cystic Fibrosis in the Genomic Era. The    Clinical biochemist. Reviews/Australian Association of Clinical    Biochemists. 2005, 26 (4), 135-153.-   8. (a) Brenner, A. J.; Harris, E. D., A quantitative test for copper    using bicinchoninic acid. Anal Biochem 1995, 226 (1), 80-4; (b)    Huang, X.; Liu, Y. H.; Chen, K. L.; Shin, W. J.; Lu, C. J.; Kong, G.    W.; Patnaik, D.; Lee, S. H.; Cortes, J. F.; Rogers, J. A.,    Stretchable, Wireless Sensors and Functional Substrates for    Epidermal Characterization of Sweat. Small 2014, 10 (15), 3083-3090.-   9. Kim, J.; Banks, A.; Cheng, H. Y.; Xie, Z. Q.; Xu, S.; Jang, K.    I.; Lee, J. W.; Liu, Z. J.; Gutruf, P.; Huang, X.; Wei, P. H.; Liu,    F.; Li, K.; Dalal, M.; Ghaffari, R.; Feng, X.; Huang, Y. G.; Gupta,    S.; Paik, U.; Rogers, J. A., Epidermal Electronics with Advanced    Capabilities in Near-Field Communication. Small 2015, 11 (8),    906-912.

Example 5: Additional Sweat Patches

FIG. 42(A) shows a schematic illustration of an epidermal microfluidicsweat sensor providing information on sweat volume and rate as well asconcentration of biomarkers in sweat incorporated with wirelesscommunication electronics and an adhesive layer for adhering the sensorto the epidermis of a subject. FIG. 42(B) shows a schematic illustrationof image process markers applied to an epidermal microfluidic sweatsensor. Image process markers are laminated on or disposed in a toplayer of the sensor for white balance and color calibration, whichenables the sensors to function under various light conditions. Theimage process markers also provide a reference for device orientationand a border-line for color change within a channel.

FIG. 43 provides graphical representations of water loss as a functionof outlet channel (A) width and (B) length. Smaller channel widthsgenerally lead to a lower rate of water vapor loss than larger channelwidths, but channel length does not significantly affect the rate ofwater vapor loss. FIG. 44 provides a graphical representation of backpressure inside a channel showing that shorter outlet channels andlarger channel widths produce lower back pressures. At a channel widthof 100 μm, back pressure became negligible for all channel lengthsstudied. The following equation was used to calculate the theoreticalpressure in the channel:

$\begin{matrix}{{\Delta \; P} \approx {\frac{8{L\left( {w + h} \right)}^{2}}{w^{3}h^{3}}\frac{\mu \; M}{\rho}\frac{{\overset{.}{V}}_{inlet}P_{0}}{RT}}} & (15)\end{matrix}$

where h=300 μm, M=29E-3 Kg/mol, ρ=1.2 Kg/m³, μ=1.8E-5 Pa·s, P₀=1E5 Pa,V_(in)=15 μL/hour, R=8.314 J/(mol·K), T=300 K.

FIG. 45 shows a schematic illustration of a cross section of amicrofluidic channel deformed due to pressure (A) and a top perspectiveview of a section of an epidermal microfluidic sweat sensor showing awidth of the microfluidic channel (B), as well as a graphicalrepresentation of deformation shown as volume change due to pressure.The volume change was calculated using:

$\begin{matrix}{\frac{\Delta \; V}{V} = \frac{\Delta Pa^{4}}{5\; {Et}^{3}h}} & (16)\end{matrix}$

where 2a=1 mm, t=100 μm, E=145 KPa and v=0.5. At an outlet width greaterthan 10 μm, a pressure-induced volume change can be avoided.

To harvest biofluids using pump-less microfluidics, sufficient adhesionforce is required to drive fluid into the microfluidics system. Thedisclosed microfluidic devices demonstrate great adhesion on theepidermis facilitated by medical-grade adhesives (e.g., Tagaderm®). FIG.46 shows an experimental set-up for 90° peel adhesion property testing(standard ISO 29862:2007) using a force gauge (Mark-10, Copiague, N.Y.)(A). A holding devices is adhered on the skin with a force gauge (B) anddevices are peeled at an angle of 90⁰ (C). The force measurement whiledisplacing the device at a rate of 300 mm/min is shown graphically in(D) and the area where peeling occurs is indicated by the gray region.The average peeling force was determined to be 5.7 N. Thus, thedisclosed microfluidic sweat sensors may be bonded to the epidermis of asubject with an adhesion force in the range from 1 N to 10 N, or 2 N to8 N, or 3 N to 6 N.

FIG. 47 illustrates one example of colorimetric determination ofcreatinine. A UV-VIS spectrum illustrating various creatinineconcentrations (i.e., 15-1000 μM) is shown in (A) and a constructedcalibration curve based on this spectrum is shown in (B). The presentedcolor for each spectrum corresponds to exhibited color on paper-basedcolorimetric detection reservoirs as a function of creatinineconcentration, which is presented in optical image (C). Thiscolorimetric analysis is based on an enzymatic reaction using a mixtureof creatinine amidohydrolase, creatine amidinohydrolase and sarcosineoxidase. Reaction of creatinine with this enzyme mixture generateshydrogen peroxide proportional to the concentration of creatinine inbiological fluids. The hydrogen peroxide concentration is determinedcolorimetrically by the chromogen 2,5-dichloro-2-hydroxybenzenesulfonicacid and 4-amino-phenazone in a reaction catalyzed by horseradishperoxidase.

FIG. 48 illustrates one example of colorimetric determination ofethanol. Ethanol is detected via reaction with alcohol dehydrogenase inthe presence of formazan dye. A UV-VIS spectrum illustrating variousethanol concentrations (i.e., 0.04-7.89% (w/v)) is shown in (A) and aconstructed calibration curve based on this spectrum is shown in (B).The presented color for each spectrum corresponds to exhibited color onpaper-based colorimetric detection reservoirs as a function of ethanolconcentration, which is presented in optical image (C).

FIG. 49 shows various microfluidic sweat sensor designs including fourindividual quantitative colorimetric detection reservoirs and anorbicular outer-circle fluidic channel. In some embodiments, a singlemicrofluidic channel is in fluidic communication with all of thecolorimetric detection reservoirs and an orbicular fluidic channel. Inan other embodiment, one microfluidic channel transports fluids from theepidermis of a subject to the colorimetric detection reservoirs and asecond microfluidic channel transports fluids from the epidermis of thesubject to the orbicular fluidic channel (C). In another embodiment,each colorimetric detection reservoir and the orbicular microfluidicchannel may be independently connected to a microfluidic channel thattransports fluid from the epidermis of a subject. Optionally, each ofthe colorimetric detection reservoirs may comprise an outlet to achannel that allows vapor to escape to the surrounding environment. Asshown in (D), the outlet channel may be tapered to increase in volumenearer the outlet to the surrounding environment, thereby accomodatinglarger quantities of vapor without increasing back pressure within theoutlet channel. In any of the embodiments disclosed, the orbicularfluidic channel may be circular or serpentine and the orbicular fluidicchannel may have a sealed distal end, optionally including a reservoir,or an outlet to the surrounding environment. As shown in FIG. 50, aserpentine orbicular fluidic channel provides a greater area and channelvolume than a circular orbicular fluidic channel while controlling forchannel width and height to avoid collapse of the channel. For example,a serpentine channel may provide an increased area of up to 58% comparedto a circular channel having an identical channel width. An increasedarea of the orbicular channel increases the amount of time amicrofluidic sweat sensor can be used for monitoring a subject withoutbeing replaced or dried.

STATEMENTS REGARDING INCORPORATION BY REFERENCE AND VARIATIONS

All references throughout this application, for example patent documentsincluding issued or granted patents or equivalents; patent applicationpublications; and non-patent literature documents or other sourcematerial; are hereby incorporated by reference herein in theirentireties, as though individually incorporated by reference, to theextent each reference is at least partially not inconsistent with thedisclosure in this application (for example, a reference that ispartially inconsistent is incorporated by reference except for thepartially inconsistent portion of the reference).

The terms and expressions which have been employed herein are used asterms of description and not of limitation, and there is no intention inthe use of such terms and expressions of excluding any equivalents ofthe features shown and described or portions thereof, but it isrecognized that various modifications are possible within the scope ofthe invention claimed. Thus, it should be understood that although thepresent invention has been specifically disclosed by preferredembodiments, exemplary embodiments and optional features, modificationand variation of the concepts herein disclosed may be resorted to bythose skilled in the art, and that such modifications and variations areconsidered to be within the scope of this invention as defined by theappended claims. The specific embodiments provided herein are examplesof useful embodiments of the present invention and it will be apparentto one skilled in the art that the present invention may be carried outusing a large number of variations of the devices, device components,methods steps set forth in the present description. As will be obviousto one of skill in the art, methods and devices useful for the presentmethods can include a large number of optional composition andprocessing elements and steps.

When a group of substituents is disclosed herein, it is understood thatall individual members of that group and all subgroups, including anyisomers, enantiomers, and diastereomers of the group members, aredisclosed separately. When a Markush group or other grouping is usedherein, all individual members of the group and all combinations andsubcombinations possible of the group are intended to be individuallyincluded in the disclosure. When a compound is described herein suchthat a particular isomer, enantiomer or diastereomer of the compound isnot specified, for example, in a formula or in a chemical name, thatdescription is intended to include each isomers and enantiomer of thecompound described individual or in any combination. Additionally,unless otherwise specified, all isotopic variants of compounds disclosedherein are intended to be encompassed by the disclosure. For example, itwill be understood that any one or more hydrogens in a moleculedisclosed can be replaced with deuterium or tritium. Isotopic variantsof a molecule are generally useful as standards in assays for themolecule and in chemical and biological research related to the moleculeor its use. Methods for making such isotopic variants are known in theart. Specific names of compounds are intended to be exemplary, as it isknown that one of ordinary skill in the art can name the same compoundsdifferently.

Many of the molecules disclosed herein contain one or more ionizablegroups [groups from which a proton can be removed (e.g., —COOH) or added(e.g., amines) or which can be quaternized (e.g., amines)]. All possibleionic forms of such molecules and salts thereof are intended to beincluded individually in the disclosure herein. With regard to salts ofthe compounds herein, one of ordinary skill in the art can select fromamong a wide variety of available counterions those that are appropriatefor preparation of salts of this invention for a given application. Inspecific applications, the selection of a given anion or cation forpreparation of a salt may result in increased or decreased solubility ofthat salt.

Every formulation or combination of components described or exemplifiedherein can be used to practice the invention, unless otherwise stated.

Whenever a range is given in the specification, for example, atemperature range, a time range, or a composition or concentrationrange, all intermediate ranges and subranges, as well as all individualvalues included in the ranges given are intended to be included in thedisclosure. It will be understood that any subranges or individualvalues in a range or subrange that are included in the descriptionherein can be excluded from the claims herein.

All patents and publications mentioned in the specification areindicative of the levels of skill of those skilled in the art to whichthe invention pertains. References cited herein are incorporated byreference herein in their entirety to indicate the state of the art asof their publication or filing date and it is intended that thisinformation can be employed herein, if needed, to exclude specificembodiments that are in the prior art. For example, when composition ofmatter are claimed, it should be understood that compounds known andavailable in the art prior to Applicant's invention, including compoundsfor which an enabling disclosure is provided in the references citedherein, are not intended to be included in the composition of matterclaims herein.

It must be noted that as used herein and in the appended claims, thesingular forms “a”, “an”, and “the” include plural reference unless thecontext clearly dictates otherwise. Thus, for example, reference to “acell” includes a plurality of such cells and equivalents thereof knownto those skilled in the art, and so forth. As well, the terms “a” (or“an”), “one or more” and “at least one” can be used interchangeablyherein. It is also to be noted that the terms “comprising”, “including”,and “having” can be used interchangeably. The expression “of any ofclaims XX-YY” (wherein XX and YY refer to claim numbers) is intended toprovide a multiple dependent claim in the alternative form, and in someembodiments is interchangeable with the expression “as in any one ofclaims XX-YY.”

Unless defined otherwise, all technical and scientific terms used hereinhave the same meanings as commonly understood by one of ordinary skillin the art to which this invention belongs. Although any methods andmaterials similar or equivalent to those described herein can be used inthe practice or testing of the present invention, the preferred methodsand materials are now described. Nothing herein is to be construed as anadmission that the invention is not entitled to antedate such disclosureby virtue of prior invention.

Every formulation or combination of components described or exemplifiedherein can be used to practice the invention, unless otherwise stated.

Whenever a range is given in the specification, for example, atemperature range, a time range, or a composition or concentrationrange, all intermediate ranges and subranges, as well as all individualvalues included in the ranges given are intended to be included in thedisclosure. As used herein, ranges specifically include the valuesprovided as endpoint values of the range. For example, a range of 1 to100 specifically includes the end point values of 1 and 100. It will beunderstood that any subranges or individual values in a range orsubrange that are included in the description herein can be excludedfrom the claims herein.

As used herein, “comprising” is synonymous with “including,”“containing,” or “characterized by,” and is inclusive or open-ended anddoes not exclude additional, unrecited elements or method steps. As usedherein, “consisting of” excludes any element, step, or ingredient notspecified in the claim element. As used herein, “consisting essentiallyof” does not exclude materials or steps that do not materially affectthe basic and novel characteristics of the claim. In each instanceherein any of the terms “comprising”, “consisting essentially of” and“consisting of” may be replaced with either of the other two terms. Theinvention illustratively described herein suitably may be practiced inthe absence of any element or elements, limitation or limitations whichis not specifically disclosed herein.

One of ordinary skill in the art will appreciate that startingmaterials, biological materials, reagents, synthetic methods,purification methods, analytical methods, assay methods, and biologicalmethods other than those specifically exemplified can be employed in thepractice of the invention without resort to undue experimentation. Allart-known functional equivalents, of any such materials and methods areintended to be included in this invention. The terms and expressionswhich have been employed are used as terms of description and not oflimitation, and there is no intention that in the use of such terms andexpressions of excluding any equivalents of the features shown anddescribed or portions thereof, but it is recognized that variousmodifications are possible within the scope of the invention claimed.Thus, it should be understood that although the present invention hasbeen specifically disclosed by preferred embodiments and optionalfeatures, modification and variation of the concepts herein disclosedmay be resorted to by those skilled in the art, and that suchmodifications and variations are considered to be within the scope ofthis invention as defined by the appended claims.

1-69. (canceled)
 70. An epidermal microfluidic sweat patch comprising: aflexible or stretchable substrate; a microfluidic channel disposed insaid flexible or stretchable substrate, the microfluidic channel havingan inlet on a bottom surface of the flexible or stretchable substrateconfigured to receive sweat released by skin sweat glands; a cavitydisposed in said flexible or stretchable substrate, wherein the cavityis fluidically connected to the inlet or the microfluidic channel forreceiving sweat released by skin sweat glands; and a plurality ofcolorimetric indicators disposed in the microfluidic channel and/or inthe cavity for multiparametric detection of a plurality of sweatparameters.
 71. The epidermal microfluidic sweat patch of claim 70,wherein said plurality of sweat parameters are two or more sweatparameters selected from the group consisting of: sweat volume and/orsweat rate; sweat pH; glucose concentration; lactate concentration;chloride concentration; creatinine concentration; copper ionconcentration; iron ion concentration and ethanol concentration.
 72. Theepidermal microfluidic sweat patch of claim 70, wherein at least one ofthe colorimetric indicators is configured to determine a sweat parameteras a function of time.
 73. The epidermal microfluidic sweat patch ofclaim 72, wherein the at least one sweat parameter as a function of timecomprises sweat volume or sweat rate.
 74. The epidermal microfluidicsweat patch of claim 73, wherein the at least one colorimetric indicatorthat determines the sweat volume or sweat rate as a function of time isdisposed in the microfluidic channel to provide an optically detectablesweat fluid leading edge in the microfluidic channel whose position canvary as a function of time.
 75. The epidermal microfluidic sweat patchof claim 70, comprising a plurality of detection reservoirs eachfluidically connected to the microfluidic channel, wherein eachdetection reservoir contains a unique colorimetric indicator.
 76. Theepidermal microfluidic sweat patch of claim 75, comprising fourdetection reservoirs for detecting four unique sweat parameters.
 77. Theepidermal microfluidic sweat patch of claim 70, wherein the stretchableor flexible substrate comprises polydimethysiloxane (PDMS) having a netbending stiffness selected such that the epidermal microfluidic sweatpatch is capable of establishing conformal contact with an epidermalsurface.
 78. The epidermal microfluidic sweat patch of claim 70, whereinthe stretchable or flexible substrate is selected to conformally contactan epidermal surface to provide a non-invasive daily-wear patch forpersonal healthcare monitoring.
 79. The epidermal microfluidic sweatpatch of claim 70, wherein the microfluidic channel is part of atwo-dimensional channel system within a PDMS substrate that directs aflow of sweat released from skin without pumps, valves or fluiddetectors.
 80. The epidermal microfluidic sweat patch of claim 79,further comprising a plurality of detection reservoirs connected to thetwo-dimensional channel system.
 81. The epidermal microfluidic sweatpatch of claim 70, wherein the microfluidic channel has a volumeselected to accommodate a volume of sweat released from skin for a timeperiod of up to 6 hours.
 82. The epidermal microfluidic sweat patch ofclaim 81, wherein the microfluidic channel comprises an orbicular outercircular or serpentine fluidic channel.
 83. The epidermal microfluidicsweat patch of claim 70, wherein the flexible or stretchable substratecomprises PDMS and has a total water loss that is less than 3% of sweatintroduced to the microfluidic channel.
 84. The epidermal microfluidicsweat patch of claim 70, wherein one colorimetric indicator comprisescobalt (II) chloride in a hydrogel matrix coated on at least a portionof a surface of the microfluidic channel for determining sweat volumeand sweat rate by optical detection of a color change in themicrofluidic channel.
 85. The epidermal microfluidic sweat patch ofclaim 84, further comprising an additional colorimetric indicatordisposed in the cavity for detection of a sweat biomarker, wherein thesweat biomarker is selected from the group consisting of: pH; glucose;lactate; chloride; copper ion; iron ion; and ethanol.
 86. The epidermalmicrofluidic sweat patch of claim 70, further comprising near fieldcommunication electronics supported by or embedded in the substrate forwireless communication with an external reader, including an externalreader that is a smartphone, for quantitative determination of theplurality of sweat parameters.
 87. The epidermal microfluidic sweatpatch of claim 70, further comprising an image process marker laminatedon or disposed in the substrate for white balance and color calibrationunder various light conditions.
 88. The epidermal microfluidic sweatpatch of claim 70, further comprising an adhesive positioned on thebottom surface of the flexible or stretchable substrate to adhere theepidermal microfluidic sweat patch to skin at an adhesion force between1 N to 10 N.
 89. A method of measuring a plurality of sweat parameters,the method comprising the steps of: applying the epidermal microfluidicsweat patch of claim 70 to a skin surface; collecting sweat releasedfrom sweat glands in the microfluidic channel via the inlet; introducingsweat to the cavity; and detecting at least two optical changes in theepidermal microfluidic sweat patch by at least two colorimetric changesof at least two colorimetric indicators to thereby measure the pluralityof sweat parameters.
 90. The method of claim 89, further comprising thestep of monitoring at least one colorimetric change over time todetermine a time course of at least one sweat parameter.